Anode Heel Effect

It is generally accepted that an x-ray beam's intensity is not uniform throughout its

entirety. As x-radiation is emitted from the target area in a conical shape,

measurements have determined that the intensity in the direction of the anode (AC) is

lower (over and above the difference caused by the Inverse Square Law) than the

intensity in the direction of the cathode (AB). The fact that the intensities vary in such

a manner causes visible differences in the density produced on the radiographs. This

phenomenon is called heel effect and is illustrated below.

NOTE:

A = 100-percent intensity

AB = consists of a slight increase over 100-percent intensity

and then a general decrease in intensity as B is approached

AC = consists of a considerable decrease in intensity as C is

approached

The decreased intensity at C results from emission which is nearly parallel to the

angled target where there is increasing absorption of the x-ray photons by the target

itself. This phenomenon is readily apparent in rotating anode tubes because they

utilize steeply angled anodes of generally 17 degrees or less. Generally, the steeper

the anode, the more severe or noticeable the heel effect becomes.

The effects of focus film distance on the visualization of heel effect are illustrated

below:

Figure G shows the film plane as having a shorter focus film distance than the film

plane in Figure H. Looking at Figure G, you can readily see that the x-ray beam's

involvement in exposing the film runs from C to B (the full cone of radiation). Heel

effect causes a greater decrease in x-ray beam intensity as one travels from the central

ray to the cathode (A to B).

As you look at Figure H, note that a long focus film distance is used which results in

the involvement of the x-ray beam at the film plane which does not utilize the full

cone of radiation (C to B). Hence, the extremities of the beam (C and B) are not used

in exposing the film. Because of this, heel effect is greatly reduced.

Anode Stator Motor

In principle, the rotating anode is the moving part of an electric motor, running in a

vacuum. The rotor carries the anode.

An external electromagnetic field, produced by a winding (stator) outside the glass

envelope, drives the rotor. Both together work as an asynchronous motor.

The air gap between rotor and stator isolates both from each other, since the winding

is electrically close to ground and the anode lays on high potential during operation.

On the other hand the gap reduces the efficiency of the rotating anode motor

significantly.

Due to this distance, the power supply for the motor must be relatively high, in order

to speed up the anode in an acceptable short time

The Rotor consists of a copper cylinder and rests in ball bearings for smooth

movement.

The bearings cannot be lubricated with ordinary grease because it would affect the

vacuum and, as a consequence, the high tension characteristics of the tube.

Soft metals such as lead and silver are applied to separate the ball bearings and the

running surfaces, in order to prevent the possibility of "jamming" in the vacuum.

This form of lubrication limits the life time of the bearings in the x-ray tube to about

1000 hours.

Therefore, the running time needs to be as short as possible, which does not allow

continuous rotation.

The rotation is controlled when a radiography is started.

The stator consist of several windings which are equally spread out around the neck

of the tube. They induce a rotating electro-magnetic field which interacts with the

rotor, causing it to rotate synchronously.

The simpliest power supply is a 220 V AC source. It was used in old generators for

the normal speed anode

A capacitor C provides the stator with a second phase.

The current in the two phases I and II have a phase shift of 120° to each other, which

produces the rotating field.

The value of the capacitor depends on the type of stator coil.

This stator is called "two-phase stator".

Diagram showing anode rotation speed from prep to expose to braking period

Anti - scatter Grid

Anti-scatter grids are simple and functional tools that improve the diagnostic quality

of radiographs by trapping the greater part of scattered radiation. Scattered radiation is

probably the biggest factor contributing to the poor diagnostic quality of radiographs.

Its effect produces a general radiographic fog on the film which reduces the contrast.

The best-known way to effectively remove the greater part of radiation scatter is

by the use of an x-ray anti-scatter grid. Radiation which does not travel in the

same direction as the primary beam is absorbed by the lead strips of the grid.

Since Dr. Gustav Bucky built the first grid in 1913, his original principle of lead

foil strips standing on edge separated by x-ray transparent interspacers has

remained one of the best-known technique to trap the scatter

Types of Filter Grids

X-ray grids are commercially available with either focused or parallel lead

strips, and these two types are produced in either linear or crossed grid

configuration. The focused grid has its leads angled progressively in such a way

that lines drawn through each lead and continued out of the gird will intersect at

a point known as the grid focus. When strips are not progressively angulated but

are perpendicular to the surface of the grid, the grid is termed "parallel" (See

Figure 1).

Both the focused or parallel grids may be made in either the linear or crossed

grid type. The linear grid is made with the lengths of all its leads in the same

direction. The crossed grid is usually two linear grids, one on top of the other,

with the leads of the top grid crossing those of the lower grid (See Figure 2).

In general, the crossed grid will remove more scattered radiation than a linear

grid of ratio equal to the combined ratios of its two parts, e.g., a crossed grid,

each of whose parts has 5:1 ratio, will remove more scattered radiation than a

linear grid of 10:1 ratio. This advantage is more striking at voltages under 100

KVP.

The advantage of the linear grid over the crossed grid is that it may be used in

tilted-tube techniques without undue "cut-off" in the radiograph. This is true

with grid ratios 8:1 and lower and only if the angle of tilt of the tube is in a

direction parallel to the length of the leads. Tilting the tube at an angle across the

leads will result in serious density reduction (cut off) on the film. With higher

ratio grids, tube angling must be slight or focal distance long to avoid marked

density variation.

Construction of Grids & Significance of Grid Ratio

The prime purpose of a grid being the absorption of stray radiation, lead strips

(the material which is most practical in the absorption of x-rays) are its basic

component. The strips -- five hundred or two thousands or more of them -- are

set on edge, properly angled to a mean focal distance and separated by x-ray

transparent interspacers. The whole is bonded together into a single flat

structure, suitably covered for strength, durability and protection against

moisture. The ratio of a grid is defined as the relation of the height of the lead

strips to the distance between them. Thus with interspacers 5 times as high as

they are wide, a grid is said to be 5:1 ratio, etc. Generally speaking, the higher

the ratio of a grid, the more scattered radiation is absorbed (see Diagram).

As grid ratio increases, the necessity of having the focused grid exactly centered

and perfectly level under the x-ray tube becomes more and more important. Also,

it becomes more necessary to use the grid as its focal distance from the tube,

instead of being able to use it through a range of distances. For example, the 40"

focal distance, 16:1 ratio grid must be used at 40" for satisfactory results, and

must be perfectly centered and leveled.

The 5:1 ratio focused grid, on the other hand, will give satisfactory results over a

wide range of focal distances, and need not to be as accurately centered or

leveled. Of course the 5:1 ratio linear grid will not have nearly the effectiveness

of secondary removal that the 16:1 has, but in ost cases this may be willingly

sacrificed to gain the latitude and ease-of-use of the low ratio grid. However, a

5:1 crossed grid will produce as good secondary removal as 16:1 grid at low kilo-

voltages, while retaining the latitude of the 5:1 ratio.

Selection Considerations

In order to prevent the shadows cast onto the film by the grid from interfering with

visualization of diagnostic detail, certain principles must be followed:

For one, the lead should be as thin as possible to be consistent with adequate absorption

of scattered radiation.The thinner the lead, the narrower the shadow it will produce on

the film and the less visible it will be to the eye. Also, the thinner it is the less absorption of primary radiation will be in the grid.

However, it must be noted that adequate absorption of scattered radiation is the function

of the grid and lead must be thick enough to provide this function. Another factor is the relative fineness of the grid. This quality is represented by the

number of lines per inch. In general, the greater the number of lines per inch, the less

visible will the individual lines be, but this is subject to certain practical considerations

which modify it in actual use.

Practical Considerations in Grid Selection:

The selection of a grid to be used for a particular radiograph will be primarily

dependent on the following considerations:

Relative quantity of scattered radiation produced by subject being radiographed.

Kilovoltage technique used.

Capacity of x-ray generator.

The quantity of scattered radiation produced is dependent on the thickness and relative density

of the body being radiographed. A non-grid exposure of the chest will consist of about one half

scattered radiation, while a non-grid exposure of the abdomen may consist of more than 90%

scattered radiation.

From this, it is apparent that for dense body sections the more effective removal of scattered

radiation will provide the most striking improvement in the radiograph. This suggests the use of

a high ratio grid or a crossed grid. The choice between these two grids depends on the ease of

aligning the grid correctly relative to the x-ray tube, and whether a high or low voltage

techniques are in use.

If there are questions about the proper centering or leveling, or if low kilovoltages are in use, a

low ratio grid will present much greater advantage from the point of view of positioning latitude

and cleanup. For high voltage techniques, if the grid can be accurately aligned (see effect of

misalignment in Figures 1 & 2 below), greater advantages will result from the use of an 8:1 ratio

crossed grid or high ratio linear grid.

At kilovoltages of the order of 100 KVP or more, comparable radiographic effect requires low

milliampere-second values than at low kilovoltages, thus reducing the radiation dosage to the

patient.

However, in order to maintain the same contrast range of the higher kilovoltage,

it is necessary to use a higher ratio grid. The exposure factors are not the same

for all ratios, and the increased exposure required for a high ratio grid may to

some extent reduce the patient-dosage advantage gained by going to higher

kilovoltage techniques. In general, in spite of the higher exposure factors

involved, the use of high kilovoltage and high ratio grids will result in somewhat

lower radiation dosage to the patient.

All radiographers must work within the limitations of the physical

characteristics of the x-ray equipment at their disposal. While this may not be as

important a consideration in the selection of a grid as some others, it is a factor

to be considered. For instance, the maximum benefits to be derived from a 16:1

ratio grid will not be realized with a unit whose top limit is 90 KVP, although

there will be some advantage over a lower ratio grid. In general, a 16:1 ratio grid

will do the most good with equipment which can be used at kilovoltages above

100 KVP.

This applies also, to a lesser extent, to the 12:1 ratio grid. With a bedside or

portable unit, where the likelihood of near-perfect alignment of the grid relative

to the primary beam is poor, the use of the high ratio grids is practically

impossible, and difficulties may be encountered even with the 8:1 ratio grids. For

such use, where wide latitude in distance, centering, and leveling is necessary, the

5:1 ratio grid is advisable, and for maximum cleanup under these conditions the

5:1 crossed grid is ideal.

Selection Guidelines

Choosing the correct grid for your application may be a difficult task. MXE provides technical

advice to assist you in selecting the proper grids and evaluating their performance.

(1) X-ray Grid Selection Based on Clean-up Requirements:

Cleanup Ratio/Type Positioning

Latitude Recommended

Up To Remarks

SUPERLATIVE 8:1 criss-cross Distance fair;

centering and

leveling-slight

120 KVP Not recommended

for tilted tube

technique

EXCELLENT 12:1 linear Very slight 110 KVP (Suitable for

highr KV) Extra care required

for proper alignment;

usually used in fixed

mount

EXCELLENT 6:1 criss-cross Good 100 KVP Tube tilt limited to

five degrees GOOD 8:1 linear Distance fair;

centering and

leveling-slight

100 KVP For general

stationary grid use

MODERATE 6:1 linear Good 80 KVP Least expensive of

stationary grids

(2) Basic Guidelines:

ANATOMY LINE RATIO DISTANCE SKULL 103 10:1 36-40" CHEST 103 10:1-12:1 60-72"

ABDOMINAL 103 8:1 34-44" SCOLIOSIS STUDIES 85-103 8:1 48-72"

SPECIAL PROCEDURES LINE RATIO DISTANCE MOST STUDIES 103 10:1 36-40"

BI-PLANE 85

criss-cross 8:1 34-44"

SURGICAL ROOM LINE RATIO DISTANCE ORTHOPEDICS 85 8:1 34-44"

CHOLANGIOGRAMS VENOUS STUDIES 103 10:1 36-40" EMERGENCY ROOM LINE RATIO DISTANCE

TRANS LATERAL SKULL, SPINES, HIPS 60-85 6:1-8:1 34-44"

Decubitus X-ray Grids

Designed to reduce grid cutoff, MXE decubitus grids position the lead strips

parallel to the short dimension of the grid-in line with the cathode-anode

direction of the x-ray tube when in the translateral position. This allows greater

positioning latitude when aligning the x-ray tube with the grid.

Difference between the standard and decubitus grid

Features of the decubitus grid:

Improved image quality-more uniform density on decubitus and BE air

contrast studies.

Ease of positioning with reduced cut-off.

Lines to short dimension recommended for use in translateral views of

skull, spine, hips ...... emergency room and surgery.

Allow portable crosswise chest radiography on large patients.

Available in a full range of sizes and ratios.

Grid Labels:

Grids are often marked with a series of idications about their properties

K is the Contrast Improvement Factor and is the ratio of the

contrast with a grid to the contrast without a grid. This factor is

dependent upon kVp, field size and thickness of tissue.

B is named after the celebrated Gustav Bucky and is the Bucky

Factor and is the ratio of incident radiation to the grid compared

with the transmitted radiation passing through the grid. It has great

practical use and is a factor that you apply when converting from a

non-grid technique to a grid technique or vice versa. The B is

dependent upon the kVp becoming larger with increased kVp.

∑ is selectivity which is usually

shown as a Sigma (like a M rotated 90 degrees anticlockwise).

This is the ratio of transmitted primary radiation to transmitted

scatter radiation and is very similar to the Primary transmission

ratio. This is a good measure of a grid because it should be high

with an efficient grid.

F is the FFD or more correctly the focus grid distance, focussed grids have an

optimum working distance

R is the Grid Ratio, the ratio of height to width of inter space material

Grid factor

Grid Factor = Exposure(mAs) with a grid

Exposure without a grid

Atomic Structure

Atoms are particles of elements, substances that could not be broken down further. In

examining atomic structure though, we have to clarify this statement. An atom cannot

be broken down further without changing the chemical nature of the substance. For

example, if you have 1 ton, 1 gram or 1 atom of oxygen, all of these units have the

same properties. We can break down the atom of oxygen into smaller particles,

however, when we do the atom looses its chemical properties. For example, if you

have 100 watches, or one watch, they all behave like watches and tell time. You can

dismantle one of the watches: take the back off, take the batteries out, peer inside and

pull things out. However, now the watch no longer behaves like a watch. So what

does an atom look like inside?

Atoms are made up of 3 types of particles electrons , protons and

neutrons . These particles have different properties. Electrons are tiny, very

light particles that have a negative electrical charge (-). Protons are much larger and

heavier than electrons and have the opposite charge, protons have a positive

charge. Neutrons are large and heavy like protons, however neutrons have no

electrical charge. Each atom is made up of a combination of these particles. Let's

look at one type of atom:

The atom above, made up of one proton and one electron, is called hydrogen (the

abbreviation for hydrogen is H). The proton and electron stay together because just

like two magnets, the opposite electrical charges attract each other. What keeps the

two from crashing into each other? The particles in an atom are not still. The

electron is constantly spinning around the center of the atom (called the nucleus). The

centrigugal force of the spinning electron keeps the two particles from coming into

contact with each other much as the earth's rotation keeps it from plunging into the

sun. Taking this into consideration, an atom of hydrogen would look like this:

A Hydrogen Atom

Keep in mind that atoms are extremely small. One hydrogen atom, for example, is

approximately 5 x 10-8

mm in diameter. To put that in perspective, this dash - is

approximately 1 mm in length, therefore it would take almost 20 million hydrogen

atoms to make a line as long as the dash. In the sub-atomic world, things often

behave a bit strangely. First of all, the electron actually spins very far from the

nucleus. If we were to draw the hydrogen atom above to scale, so that the proton

were the size depicted above, the electron would actually be spinning approximately

0.5 km (or about a quarter of a mile) away from the nucleus. In other words, if the

proton was the size depicted above, the whole atom would be about the size of Giants

Stadium. Another peculiarity of this tiny world is the particles themselves. Protons

and neutrons behave like small particles, sort of like tiny billiard balls. The electron

however, has some of the properties of a wave. In other words, the electron is more

similar to a beam of light than it is to a billiard ball. Thus to represent it as a small

particle spinning around a nucleus is slightly misleading. In actuality, the electron is a

wave that surrounds the nucleus of an atom like a cloud. While this is difficult to

imagine, the figure below may help you picture what this might look like:

Hydrogen: a proton surrounded by an electron cloud

While you should keep in mind that electrons actually form clouds around their

nucleii, we will continue to represent the electron as a spinning particle to keep things

simple.

In an electrically neutral atom, the positively charged protons are always

balanced by an equal number of negatively charged electrons. As we have seen,

hydrogen is the simplest atom with only one proton and one electron. Helium is the

2nd simplest atom. It has two protons in its nucleus and two electrons spinning

around the nucleus. With helium though, we have to introduce another

particle. Because the 2 protons in the nucleus have the same charge on them, they

would tend to repel each other, and the nucleus would fall apart. To keep the nucleus

from pushing apart, helium has two neutrons in its nucleus. Neutrons have no

electrical charge on them and act as a sort of nuclear glue, holding the protons, and

thus the nucleus, together.

A Helium Atom

As you can see, helium is larger than hydrogen. As you add electrons, protons

and neutrons, the size of the atom increases. We can measure an atom's size in two

ways: using the atomic number (Z) or using the atomic mass (A, also known as the

mass number). The atomic number describes the number of protons in an atom. For

hydrogen the atomic number, Z, is equal to 1. For helium Z = 2. Since the number of

protons equals the number of electrons in the neutral atom, Z also tells you the

number of electrons in the atom. The atomic mass tells you the number of protons

plus neutrons in an atom. Therefore, the atomic mass, A, of hydrogen is 1. For

helium A = 4.

Ions and Isotopes So far we have only talked about electrically neutral atoms, atoms with no

positive or negative charge on them. Atoms, however, can have electrical

charges. Some atoms can either gain or lose electrons (the number of protons never

changes in an atom). If an atom gains electrons, the atom becomes negatively

charged. If the atom loses electrons, the atom becomes positively charged (because

the number of positively charged protons will exceed the number of electrons). An

atom that carries an electrical charge is called an ion. Listed below are three forms of

hydrogen; 2 ions and the electrically neutral form.

H+ : a positively charged

hydrogen ion

H : the hydrogen

atom

H- : a negatively charged

hydrogen ion

Neither the number of protons nor neutrons changes in any of these ions,

therefore both the atomic number and the atomic mass remain the same. While the

number of protons for a given atom never changes, the number of neutrons can

change. Two atoms with different numbers of neutrons are called isotopes. For

example, an isotope of hydrogen exists in which the atom contains 1 neutron

(commonly called deuterium). Since the atomic mass is the number of protons plus

neutrons, two isotopes of an element will have different atomic masses (however the

atomic number, Z, will remain the same).

Two isotopes of hydrogen

Hydrogen

Atomic Mass = 1

Atomic Number = 1

Deuterium

Atomic Mass = 2

Atomic Number = 1

Attenuation of X-rays

The percentage of X-ray energy absorbed by the material is due to a process known as

electron ionisation , this is dependent upon the material density and atomic number.

As a result the detected X-ray attenuation provides a picture of the absorbed energy

on the irradiated objects. Due to the absorbed energy being relative to the atomic

number, it can be used in the material discrimination process.

Generally the lower the atomic number the more transparent the material is to the X-

rays. Materials composed of elements with a high atomic numbers absorb radiation

more effectively causing less dark shadows in an X-ray image. Substances with low

atomic numbers absorb less X-ray radiation, hence their shadowgraph appears a

darker colour.

The absorption of the X-ray radiation by a material is proportional to the degree of X-

ray attenuation and is dependent on the energy of the X-ray radiation and the

following material parameters:

thickness;

density;

atomic number;

The attenuation or absorption, usually defined as the linear absorption coefficient, µ,

is defined for a narrow well-collimated, monochromatic x-ray beam. The linear

absorption coefficient is the sum of contributions of types of attenuation as listed

below.

Mass attenuation coefficient is defined as the linear attenuation coefficient divided

by the density of the medium. For a given incident gamma ray energy, the mass

attenuation coefficient is independent of the physical and chemical state of the

absorber. Thus, the mass attenuation coefficient is the same for water whether present

in liquid or vapor form

Interactions of X-Rays with Matter

The dependence of the X-ray attenuation on the atomic number relies on mainly on

three phenomena: photoelectric effect, Compton effect and pair production;

The photoelectric effect is predominant at low X-ray energies and with high atomic

numbers. When a quantum of radiation strikes an atom, it may impinge on an electron

within an inner shell and eject it from the atom. If the photon carries more energy than

is necessary to eject the electron, it will transfer this residual energy to the ejected

electron in the form of kinetic energy

The Compton effect occurs primarily in the absorption of high X-ray energy and low

atomic numbers. The effect takes place when high X-ray energy photons collide with

an electron. Both particles may be deflected at an angle to the direction of the path of

the incident X-ray. The incident photon having delivered some of its energy to the

electron emerges with a longer wavelength. These deflections, accompanied by a

charge of wavelength are known as Compton scattering.

Pair production is the formation or materialization of two electrons, one negative

and the other positive (positron), from a pulse of electromagnetic energy traveling

through matter, usually in the vicinity of an atomic nucleus. Pair production is a direct

conversion of radiant energy to matter. It is one of the principal ways in which high-

energy gamma rays are absorbed in matter. For pair production to occur, the

electromagnetic energy, in a discrete quantity called a photon, must be at least

equivalent to the mass of two electrons. The mass m of a single electron is equivalent

to 0.51 million electron volts (MeV) of energy E as calculated from the equation

formulated by Albert Einstein, E = mc2, in which c is a constant equal to the velocity

of light. To produce two electrons, therefore, the photon energy must be at least 1.02

MeV. Photon energy in excess of this amount, when pair production occurs, is

converted into motion of the electron-positron pair. If pair production occurs in a

track detector, such as a cloud chamber, to which a magnetic field is properly applied,

the electron and the positron curve away from the point of formation in opposite

directions in arcs of equal curvature. In this way pair production was first detected

(1933). The positron that is formed quickly disappears by reconversion into photons

in the process of annihilation with another electron in matter.

Two less important (In diagnostic energy levels) effects

Thomson scattering (R), also known as Rayleigh, coherent, or classical scattering,

occurs when the x-ray photon interacts with the whole atom so that the photon is

scattered with no change in internal energy to the scattering atom, nor to the x-ray

photon. Thomson scattering is never more than a minor contributor to the absorption

coefficient. The scattering occurs without the loss of energy. Scattering is mainly in

the forward direction.

Photodisintegration (PD) is the process by which the x-ray photon is captured by the

nucleus of the atom with the ejection of a particle from the nucleus when all the

energy of the x-ray is given to the nucleus. Because of the enormously high energies

involved, this process may be neglected for the energies of x-rays used in radiography.

Absorption Edges

If the mass absorption coefficient of a material is plotted against wavelength as shown

in Figure Y for a monochromatic x-ray beam, m shows sharp discontinuities at

particular wavelengths.

Fig Y

These correspond to the ionisation energy of a K shell electron and indicate the

increased probability of photoelectric absorption, however this drops sharply as the

difference between the photon and electron binding energy increases. The variation of

m with photon energy E and atomic number Z for the various scattering and

absorption processes is summarised in the following table and shown graphically in

figure X:

Summary of Main Attenuation Mechanisms

Mechanism Variation of m with E Variation of m with Z Energy range in

tissue

Rayleigh 1 / E Z2 1 - 30 keV

photoelectric 1 / E3 Z

3 1 - 100 keV

Compton falls gradually with E independent 0.5 - 5 MeV

pair production rises slowly with E Z2 > 5 MeV

The relative Importance of Attenuation processes

Only photoelectric effect and Compton effect are significant in the production of

diagnostic radiographic images

Figure X

http://img.cryst.bbk.ac.uk/www/kelly/medicalxrays.htm

Automatic Exposure Control (AEC)1

Basic X-Ray Diagram

(Siemens)

Automatic exposure control using an ionisation chamber between the patient and the

film cassette is used in the majority of x-ray generators. A slim panel containing three

ionisation chambers (dose detector) is placed between the grid and the cassette in the

bucky assembly. The chamber assembly is constructed to be of very low density so as

not to interfere with the image.

The principal of operation is that during an exposure the air in the chamber/s is

ionised permitting a current o flow through them, this current is used to charge a

capacitor. When the capacitor voltage reaches a pre determined level the voltage is

used to terminate the exposure via a thyristor based additional control unit.

Film Screen Control

The system is set up during installation to work with the type of film screen

combinations used in the department, the setup is designed to work with the film

screen combination and the Kv selected to ensure a consistent film density with

various film screen combinations and Kv choices.

Minor density control There is a small resistor in series with the capacitor to provide the operator with a

small degree of control of the resultant film density, to allow for the patient build and

the amount of scatter produced by the radiographic technique.

Ionization chamber patterns

The unit manufacturer should provide a Perspex slide to fit the LBD to show where

the chambers are sited, in general there are three chambers a central chamber and two

outer chambers, care must be taken to ensure that the patient part being imaged lies

under the selected chamber or chambers. The chambers are often oblong in at centre

and round on the outer ones and about 7cm x 5cm and 5 cm round.

Control Panel Controls

Chamber Selection

AEC in use

Care must be taken in the use of AEC devices to ensure that the a chamber is in the

field of radiation when the exposure is made. It is also important to ensure the correct

chamber/s selection is made and that the filed size is restricted to a minimum to

reduce scatter but must cover the chamber.

The object density and thickness must be such that the minimum response time of the

system is allowed for, using high tube currents may produce exposure times shorter

than the equipment can reliably cope with. There is usually some indication on the

control desk when an error occurs such as too short a time or the generator cannot

give enough exposure, there is usually a facility for reading the mAS given during the

exposure.

AEC and tomography

There are some machines which have AEC on tomography these setups are different

as it is not possible to reliably adjust the filament current during exposure, the

exposure control is performed on the KV.

AUTOMATIC EXPOSURE CONTROL2

I. INTRODUCTION

Automatic exposure control devices can assist the radiographer in producing consistent radiographic images from patient to patient, regardless of size or presence of pathology. The advantages of this consistency are numerous and include: decreased repeat rate; decreased patient exposure; and increased department efficiency. The most important benefit being decreased repeat rate. According to Chesney's, Equipment for Student Radiographers, "Surveys conducted on a wide scale have drawn conclusions that inaccurate exposures have been the most common cause for radiographs needing to be repeated" (1994).

Although automatic timers have the potential to decrease the amount of films and increase department efficiency, this can only be accomplished if the equipment is operated by a skillful technologist. Even though it is called an "automatic" exposure device, a technologist must be very knowledable about automatic timing devices to produce high quality radiographs.

This course is designed to review the basic operation of Automatic Exposure Control devices and offer suggestions to the participant on how to utilize AEC devices to obtain optimum radiographs.

1.1 METHODS OF TERMINATING AN EXPOSURE

There are two ways that a radiographic exposure can be terminated: manually or automatically. When an examination is manually timed, the technologist sets the kVp, mA, and time. After the predetermined time has elapsed, the exposure is terminated. If the equipment is operating properly and the correct technique was used for the appropriate patient thickness, one can expect a properly exposed radiograph.

When an AEC device is used to terminate an exposure, the technologist sets the kVp and mA, but the time of the exposure is automatically determined by the machine. The AEC device differs from a manual timer because the AEC does not stop the exposure until the film has reached an appropriate density.

Unlike manual timers, which simply stop the exposure after the preset time has elapsed.

A major benefit of the AEC device is its ability to consistently obtain accurately exposed radiographs, even in the presence of pathology. While manual timers terminate the exposure at the preset time, regardless of pathology or achievement of proper film density. The following example demonstrates the difference:

Two patients may come to the radiology department for chest x-rays. They both may measure 18 centimeters, one may have normal lung fields, while the other may have a pleural effusion. Since both patients measure the same thickness, the radiographer would most likely use the same technique on both patients when manually timing.

Chances are the radiograph of the fluid-filled lungs will be lighter than the healthy lungs and therefore it would have to be repeated. If AEC would have been used in this situation, the exposure time would have been automatically increased to compensate for the fluid in the lungs . A diagnostic radiograph would have been produced, therefore eliminating the need for a repeat radiograph.

II. AEC PHYSICS AND INSTRUMENTATION

AEC devices are common in today's radiographic equipment. When AEC devices were first introduced, they were strictly used for fluoroscopic spot films. As advances were made in technology, automatic timers were re-designed to be used in wall and table buckys. Today, AEC devices find application in general, fluorosopic and even portable radiographic equipment. Before discussing AEC devices any further, it is important to review the basic operation of an automatic timer.

2.1 TYPES OF AUTOMATIC EXPOSURE CONTROLS

The most common type of AEC devices used in today's radiographic equipment is the ionization chamber. In older equipment, the phototimer was most commonly used for the automatic timing mechanism.

Even though the ionization chamber and the phototimer operate differently, they both have the same function: convert radiation

into an electrical signal which will be used to automatically stop the exposure when the film has reached the proper density.

2.2 THE IONIZATION TIMER

The ionization timer utilizes an ionization chamber, capacitor, and exposure terminating switch to automatically terminate the exposure after the film has reached the predetermined density. A brief review of how the ionization timer operates, will be beneficial at this point.

An ionization chamber is a radiation detection device that produces a small electrical current when struck by radiation. Inside the chamber are two conducting plates which are separated by air. When radiation strikes the chamber, the air inside the chamber is ionized and the electrons migrate toward the plates, thus producing a small electrical current. This electrical current is then used to charge a capacitor. When the capacitor, an electronic storage device, reaches a predetermined charge the exposure terminating switch is activated and the radiographic exposure is terminated. It is important to remember that in this situation, the exposure was terminated by sensing the amount of radiation reaching the film, and not by a preset time.

When an x-ray machine utilizes the ionization chamber as its automatic timer, three chambers are used in the configuration that is demonstrated in Figure 2.2. The ionization chambers are usually located behind the grid and in front of the cassette.

Figure 2.2

2.3 THE PHOTOTIMER

Another type of automatic exposure device that may be used in radiographic equipment is the phototimer. The phototimer was commonly used in older x-

ray equipment and consisted of a fluorescent screen, photomultiplier tube, capacitor and an exposure terminating switch. Although the components of the phototimer are different, the theory of operation is similar to that of the ionization timer which is discussed in section 2.3.

The ionization timer and phototimer both convert radiation into an electrical signal which is used to terminate the exposure when the film has reached the proper density. However, since the phototimer is a bit less sophisticated than the ionization chamber, there are a few more steps involved in the conversion of radiation into an electrical signal.

In the phototimer assembly, a fluorescent screen is placed behind the bucky. A photomultiplier tube is then placed directly behind the fluorescent screen. The fluorescent screen converts the radiation that exits from the patient and cassette into light. The photomultiplier tube then converts the light emitted from the screen into an electrical current which is used to charge a capacitor. When the capacitor is charged to the predetermined level, the exposure is terminated.

As with the ionization timer, the length of the exposure is based on the time it takes to charge the capacitor to the predetermined level and not a time set by the technologist. If a patient is larger than "average" more radiation is absorbed by the patient, and less is converted to an electrical signal. Therefore, the exposure will be longer since it will take longer to charge the capacitor to the predetermined level. Likewise, if a patient is very thin, there is minimal absorption of the beam which results in more radiation being converted into an electrical signal. This in turn will charge the capacitor more quickly and terminate the exposure more rapidly. Since the xray tube is "on " while the capacitor is being charged, it should become obvious that the length of time that it takes to charge the capacitor to the predetermined level is directly related to film density.

As mentioned earlier, the ionization timer is used more commonly in modern radiographic equipment than the phototimer. However, the term "phototiming" has become synonymous with either type of automatic exposure control.

Now that the basic operation of AEC devices have been reviewed, it is time to discuss how to properly use them.

III. TECHNOLOGIST'S DECISIONS

When using an automatic exposure control device, there are many important decisions a technologist must make in order to ensure that a diagnostic film will be obtained. The two most important are patient positioning and proper detector selection. As discussed earlier, the Phototimer or Ionization Chamber which are known as detectors collect the radiation coming from the patient and convert it into an electrical signal. For the film to have the proper density, the detector must sample the radiation coming directly from the area of interest. If the detector samples radiation from another area, the film will not

have the proper density. This then explains why proper positioning is so important when when using AEC devices: Incorrect positioning will lead to a film with incorrect density.

3.1 PATIENT POSITIONING

The following diagrams will help illustrate the importance of proper patient position.

In the diagram of an incorrectly positioned shoulder, the radiograph will not have the proper density because the shoulder joint is not directly over the detector. Due to the poor positioning, a portion of the detector is completely outside of the body and will be directly exposed by the beam. This will charge the capacitor very quickly, resulting in a radiograph that is too light to fully demonstrate the shoulder.

In Figure 3.1-1, the lateral spine radiograph will be too light because the detector is sampling radiation from soft tissue along with radiation emerging from the spine. Because the soft tissue is easily penetrated, a "large" electric current will be produced in the timing circuit. The capacitor will be charged quickly resulting in a radiograph that is too light because the exposure time was too short.

Figure 3.1-1

Without a technologist who is very knowledgeable about anatomy and positioning, automatic timers are worthless. In fact, they may actually decrease department efficiency because of the increased amount of repeat radiographs that will result if used improperly.

3.2 DETECTOR SELECTION

Along with proper positioning, proper detector selection also influences the operation of the AEC device. Auto timers may have one to three detectors in their circuits, most table and wall buckys have three. Because there are three, the question often arises, "Which detector should I use?" If one keeps in mind that the detector must sample radiation coming from the area of interest, the decision of which detector to select becomes an easy one.

Here are some general guidelines to follow for determining detector selection:

When the vertebral column is the main area of interest, the center detector should be selected.

When using AEC for joints such as the shoulder or knee, the center detector should be selected. If the outside detectors were selected, the radiograph may not have the proper density because the outer detectors may be collimated out of the field or they may detect too much radiation coming from the soft tissue. Either one of these situations will result in a radiograph that does not have the proper density.

When the pelvis is being radiographed using AEC, the two outside detectors should be selected. When a pelvis is properly positioned the two outer detectors will be directly below the ilia.

Detector selection for a chest x-ray is a bit more challenging because several factors influence this decision. Radiologist's preference, pathology, and surgical intervention play the biggest role in choosing which detector to use.

When the lungs are the area of interest, the right or both outer cells may be selected. The use of both outer cells for a PA chest radiograph will result in a slightly darker radiograph, since the left cell will take longer to accumulate

radiation due to absorption by the heart. Therefore, when deciding between right or both outer cells for a PA chest radiograph, one should consider if the radiologist prefers darker or lighter chest films.

As a general rule, use of the left or center detector for a PA chest radiograph will result in an over exposed radiograph except in the presence of certain chest pathology or surgical intervention.

If a large pleural effusion is present in a lung and the detector over the affected lung is selected, an overexposed radiograph will result. This is explained by the fact that the fluid in the affected lung will absorb a greater amount of radiation, which in turn will result in less radiation getting to the detector. Since less radiation reaches the detector, the exposure will continue longer, and an overexposed film will result. In this situation, the cell opposite the affected lung should be selected.

If a patient has undergone a pneumonectomy, the detector on the unaffected side should be selected when using AEC for the PA chest radiograph. If the cell under the affected side was selected in this situation, an underexposed radiograph would result. Since the side of the surgical intervention would offer little absorption of the radiation, the detector would accumulate radiation very rapidly resulting in a short exposure. The film would be undiagnostic because the remaining left lung and mediastinum would not be visualized adequately.

The center cell should be selected if the mediastinum is of interest on the PA chest radiograph, and also for the lateral chest xray.

When using AEC for the abdomen, the technologist's choice for detector selection is once again related to the radiologist's preference for darker or lighter radiographs. Although the center, outer two, or all three detectors maybe used for the KUB; selection of the outer two cells is most technically accurate. The KUB radiograph is most commonly ordered to evaluate the soft tissue structures of the abdomen. Selection of the two outer detectors will sample the radiation coming from the soft tissue structures only, resulting in a properly exposed radiograph for the area of interest. Selection of the center cell will result in a slightly darker radiograph because the lumbar spine will attenuate a greater portion of radiation compared to soft tissue, therefore resulting in a darker radiograph. Finally, use of all three detectors will result in a radiograph having density midway between a radiograph taken with the outer detectors and a radiograph taken using only the center detector. This is explained by the fact that the detectors are sampling a portion of the radiation coming from the soft tissue and bony structures. Therefore, an electronic

averaging occurs between those structures.

Figure 3.2-1 demonstrates the effect of proper cell selection and its affect on density. Radiograph A was taken with the the two outer cells selected, while radiograph B with take with the center cell selected.

Figure 3.2-1

When using AEC for an upright abdomen, the center detector should be selected. Use of the outer two or all three detectors is not a good choice because of the configuration of the detectors. Since the outer two detectors are positioned higher and more laterally than the center detector, there is a chance these detectors may sample radiation coming from the base of the lungs. Because the lungs are easily penetrated, the radiographic exposure will be terminated prematurely, resulting in a radiograph that is too light. Figure 3.2-2 demonstrates that when using the outer cells for an upright abdomen, the film will lack sufficient density.

figure 3.2-2

* It should be noted that these suggestions for cell selection are based on operational theory. Due to differences in equipment and calibration, results

may vary from machine to machine.

Cassette size is another important factor that should be considered when choosing which detector to use for an exposure. When using cassette sizes that are smaller than 10" x 12", only the center detector should be used. Cassette sizes smaller than 10" x12 " have such a small area of coverage that a portion of the two outer detectors lie outside of the collimated area. If the outer detectors where selected in this situation, a portion of the radiation

would never be able to reach the entire detector therefore resulting in a longer time to charge the capacitor. This would result in an unnecessarily longer exposure time and an overexposed film.

A thorough understanding of the AEC makes detector selection less threatening. Simply remember the AEC must sample radiation coming from the area of interest.

3.3 DENSITY SELECTOR

Along with proper positioning skills and knowledge of detector selection, the radiologic technologist must understand the function of the density selector to fully utilize AEC.

When the automatic timer is installed, the capacitor in the circuit is set to terminate the exposure when it has acquired a specific charge. Remember, that as long as the capacitor is charging, the x-ray tube is producing radiation. The precise charge on the capacitor which terminates the exposure is determined by phantom studies conducted by the service engineer and radiologist input. The service engineer then adjusts the Neutral setting of the density control to correspond to an acceptable radiographic density. When a radiographic imaging unit is properly calibrated, the density selector should be kept on the "neutral" setting for the majority of procedures. If you find that you are not using the "neutral" setting for most of the exams, the unit most likely needs to be recalibrated.

There are occasions when the radiographic machine is properly calibrated, but using the neutral setting does not result in properly exposed radiographs. Some of the reasons for this will now be discussed.

Probably the biggest culprit of improper film density while using AEC is inaccurate positioning. As mentioned earlier, if the part is not properly positioned over the detector, the film will not have the correct density. The density control cannot be blamed for inadequate film density if the patient is not properly positioned!

In figure 3.3-1, the detectors have been outlined with lead wire. These radiographs demonstrate how improper positioning affects density.

Radiograph B is properly positioned and therefore has sufficient density. In radiograph A, the central ray is directed too far posteriorly and in radiograph C, the central ray is directed too far anteriorly.

A B C

figure 3.3-1

Poor Collimation can be another cause for obtaining improperly exposed radiographs while using the AEC The detectors cannot tell the difference between the primary radiation coming from the patient and scattered radiation. Therefore, if large amounts of scattered radiation are being produced, it will be picked up by the detector(s) and cause the exposure to terminate too soon, resulting in a radiograph that has insufficient density. This radiograph will be too light, even though it was taken with density selector set at Neutral.

The radiographs in Fig. 3.3-2 demonstrate that as you increase the amount of collimation and shielding, the amount of scattered radiation reaching the detector decreases, resulting in a darker diagnostic radiograph.

Figure3.3-2

Although, little thought about, another factor that can affect the density of a radiograph while using AEC is use of the proper film/screen combination. When the automatic timer is calibrated, it is adjusted so the Neutral setting will produce a properly exposed film for a specific film/screen combination. If a cassette with a different film/screen combination is used, the film will not have the correct density because the AEC device cannot recognize changes in film/screen speed. Figure 3.3-3 demonstrates how film/screen speeds affect density.

Figure 3.3-3

The AEC device was calibrated to be used with the film screen combination demonstrated on the left. The film on the right has a film/screen speed

combination that is sixteen times slower (extremity) than the film on the left. Both radiographs were taken at the "N" setting.

Figure 3.3-4 is another example of how film/screen combinations affect density. The machine was calibrated to be used with the film/screen combination of the radiograph on the left. The film on the right was taken with a faster film/screen speed combination, resulting in a darker film.

Figure 3.3-4

When using AEC devices, it is important to remember to use only the film/screen combinations that were calibrated to be used with the machine.

It is evident that the previously discussed causes of improper film density can be attributed to "user error". All too frequently AEC devices "take the rap" for improperly exposed radiographs. More often than not, AEC errors are most likely the result of poor positioning and/or collimation. However, there are situations where the technologist performed the procedure correctly but ended up with an improperly exposed film when using an AEC device at the neutral density setting. Certain patient conditions such as obesity, and ascites may

affect the operation of the AEC. Because of excessive scatter produced in these situations, the AEC device may terminate the exposure prematurely.

Surgical intervention may also have an affect on the ability of the AEC device to operate properly. As discussed earlier, if a patient has had a pneumonectomy and the detector that is selected is on the same side as the pneumonectomy the image will not have the proper density.

Another factor that may influence the performance of the AEC is the presence of a prosthesis. Should the detector be directly behind a prosthesis, the radiograph will be too dark. This is easily explained by the fact that the prosthesis will prevent radiation from reaching the detector, which in turn increases the length of the radiographic exposure.

Finally, one must also consider machine failure if films taken using an AEC device do not have the proper density.

3.3.1 CHANGING THE DENSITY SETTINGS

When a radiograph needs to be repeated because of improper density, and it is not due to user error such as incorrect positioning or improper collimation, the density selector switch must be used to change the density. Since the density of a radiograph taken using AEC is determined by how long it takes to charge a capacitor to the predetermined level, the mA and time selectors no longer control the density of the radiograph as it did in manual timing. Basically, when the density selector switch is changed, the resistance in the timer circuit is also changed. Increasing the density selector by using +1 or +2 increases the resistance in the circuit, which in turn causes the capacitor to take longer to be charged to its predetermined level. Therefore, since it will take longer to charge the capacitor, the radiograph will have greater density. Likewise, choosing -1 or -2 on the density selector, deceases the resistance in the timer circuit which allows the capacitor to be charged faster than normal. This results in a film having decreased radiographic density.

At this point, it would be beneficial to review how the characteristics of kVp, mA and time change when using AEC, compared to manual timing. This will assist in understanding how to make corrections for improperly exposed radiographs obtained while using AEC devices.

As with manual timing, kVp still controls contrast when using AEC devices. However, the practice does exist where technologists increase kVp to increase density when an underexposed film is obtained while using AEC. This is an incorrect solution to the problem since kVp primarily controls contrast. A repeat film taken with higher kVp may appear darker than the original, however that is due to the decreased contrast of the film.

Figure 3.3.1-1 demonstrates that as you increase kVp, the films do appear darker, but as mentioned above, it is result of the change in the scale of contrast.

Figure 3.3.1-1

Increasing kVp to produce a darker film should be avoided because the resultant change in the scale of contrast may affect interpretation of the radiograph.

During the use of the AEC device, the function of the mA control changes dramatically. No longer does the mA setting govern density, as it did while using manual timing. When using an automatic timer, the mA control now influences the time of the exposure. The mA selector still controls the quantity of radiation produced, however when using AEC one must remember the length of exposure is determined by the time is takes to charge the capacitor to a specified level. Therefore when using AEC devices, milliamperage influences the time of exposure because when a higher mA is selected more radiation is present. When more radiation is present, the capacitor is charged more rapidly resulting in a shorter exposure.

The timer control becomes inoperable when using AEC. This makes sense because the main purpose of the AEC is to terminate the exposure automatically. Certain manufacturers design their equipment so that when AEC is selected, the timer control becomes the backup timer control. This will be discussed in greater detail later in this unit.

Based on the above information, it should become obvious that when using the AEC device, density changes should only be made by using the density selector.

The density selector allows the technologist to increase or decrease the density in predetermined increments. It is a good idea to know how each density step affects the density of the radiograph so logical predictions can be made of which setting to use should a repeat be necessary. Most radiographic equipment is calibrated to that +1 corresponds to a 25% increase in density from the "N" setting and +2 corresponds to a 50% increase. The -1 setting corresponds to 25% decrease in density compared to the "N" setting, while the -2 setting corresponds to a 50% decrease. Be aware that the density settings may be adjusted for any desired increments by the service engineer.

Figure 3.3.1-2 demonstrates how density selector changes directly affect film density. The density settings are listed below each knee image. Evaluation of these images with a densitometer did indicate the following:

A 50% decrease in density from Radiograph C (Neutral)

B 25% decrease in density

D 25% increase in density

E 50% increase in density

A B C D E

-2 -1 N +1 +2

IV. MINIMUM RESPONSE TIME

Occasionally a technologist may obtain an overexposed film on the -2 density setting. Even if the technologist positioned the patient properly and correctly collimated, there are times when the radiograph may still be too dark. This situation is due to a property of automatic timers known as minimum response time.

After the capacitor in the automatic timer circuit has been charged to the predetermined level, a signal is sent to terminate the exposure. Unfortunately, the exposure is not terminated instantaneously. Anywhere from three to thirty milliseconds may elapse before all the electronics and relays actually stop the exposure. This "lag" time is referred to as Minimum Response Time (MRT). MRT poses a problem to the radiographer when the MRT is longer than the

time required for the exposure. In situations where the MRT is longer than the required exposure time, the radiograph will be too dark regardless of the density setting. An example will help clarify this concept:

The minimum response time of a machine is 5 milliseconds. An elderly woman with emphysema requires a chest xray. As a result of her condition, the woman's chest x-ray will only require a 2 millisecond exposure. Even though the required time for the exposure is only 2 milliseconds, the actual exposure time will be 5 milliseconds because it takes the machine a minimum

of 5 milliseconds to terminate an exposure. Therefore, the radiograph will be dark when using the "N", "-1" or "-2" settings of the density selector because the machine simply cannot shut off in two milliseconds.

Problems with MRT are more common in older equipment since the MRT of these units can be as long as 30 milliseconds. Newer equipment is not faced with MRT problems as frequently as older units because advances in technology have reduced MRT's to as little as 1 millisecond.

4.1 SOLUTIONS

Problems associated with long MRT's can be solved by the use of a lower mA setting, decreased kVp, or simply using manual timing. The use of a lower mA setting is the preferred method to correct MRT problems.

When using a lower mA setting, the required time for the exposure is increased (less radiation is present, therefore a longer exposure). The goal is to make the present exposure time longer than the MRT of the machine. By making the the exposure time longer than the MRT, the automatic timer can now terminate the exposure at the proper time which in turn will result in a film with the correct density.

If the mA cannot be selected while using AEC (i.e. portable units equipped with AEC), the kVp can be lowered which will lead to an increased exposure time. Lower kVp settings result in decreased x-ray production, therefore it will take longer to charge the capacitor to appropriate level resulting in a longer exposure. However one must be careful when using kVp to change exposure time because of the affect on contrast.

Finally,the technologist should always consider one definite way to solve problems associated with the MRT of automatic timer and that is to use manual timing.

V. THE BACK-UP TIMER

One other feature that must addressed when discussing AEC devices is the back-up timer. In order to minimize errors which lead to repeat radiographs, the technologist must be fully aware of the purpose and operation of the back-up timer.

The back-up timer is a safety device which prevents the patient from receiving an excessive dose of radiation should the automatic timer fail due to mechanical or operator errors.

The back-up timer automatically terminates a phototimed exposure if it exceeds 600mAs (Federally mandated). Some radiographic units allow the operator to set the back-up time, while others are fixed at 600 mAs. If the radiography machine allows the back-up time to be adjusted, a good rule of thumb to follow is to set the back-up timer for two to three times the estimated

mAs of the exposure. It is poor practice to always set the back-up timer to the maximum level because if the automatic timer fails, the patient will be unnecessarily exposed to the radiation produced by a 600 mAs exposure before radiation production is actually terminated. If the machine allows the user to select the back-up time for the exposure, the mA selector and timer control automatically become the back-up timer selector.

5.1 REASONS WHY THE BACKUP TIMER IS ACTIVATED

The are several reasons why the back-up timer may be activated and the exposure terminated. Excessively large patients or pathologic conditions may activate the back-up timer. However, it is reasonably safe to say that the main reason the back-up timer is activated during an exposure is a careless technologist. At this point, it will be helpful to review the most common errors that cause the back-up timer to be activated:

Wrong Detector Selected: If a patient is being radiographed on the table and the chest board detectors are accidently selected, the film will be too dark and the back-up timer will be activated. This occurs because the selected detectors on the chest board did not receive any radiation so the exposure continued until it was terminated by the back-up timer. Unfortunately, should this occur, the patient is exposed by the primary beam much longer than is necessary.

Wrong Tube Selected: If a procedure room has two radiographic x-ray tubes and the tube over the bucky is not the one that is energized, the exposure will continue until it is terminated by the back-up timer. This is explained by the fact that tube over the selected detector is not producing any radiation, therefore the exposure would continue indefinitely if not terminated by the back-up timer.

Incorrect Tube/Bucky Alignment: If the x-ray tube is not centered to the bucky, the back-up timer may be activated. Since the detectors are located in the bucky assembly, they cannot sense any radiation if they are not aligned with the tube. Therefore, the exposure will continue longer than necessary.

Incorrect Back-up Timer Setting: If the back-up timer is set for too short of a time, the back-up timer will terminate the exposure prematurely, resulting in an underexposed radiograph. This happens most frequently if the previous exam used a short manual time and the back-up timer was not set before the exposure was made.

An important point to remember about the back-up timer is that it is a safety feature which is used to prevent the patient from receiving excessive radiation due to machine failure. More often than not, the back-up timer is activated because of technologist error instead of equipment failure. Unfortunately, when the back-up timer is activated the patient has already been improperly exposed to radiation, and a repeat radiograph will be necessary resulting in additional unnecessary radiation exposure to the patient. Therefore, a

radiologic technologist should always make sure all the automatic timer controls are properly set before the radiographic exposure is made.

VI. SUMMARY

The development of automatic exposure control has brought with it the possibility of decreased repeat rates and increased productivity. However, this is only possible with a technologist who fully understands all of the steps that are necessary to produce a diagnostic film while using AEC. Technologists who do not have a complete understanding of AEC may actually find themselves with an increased repeat rate, since its use requires precise positioning and thorough knowledge of the equipment. Perhaps a better term for automatic exposure control should be "assisted" exposure control because in reality the technologist does assist the automatic timing device in producing a diagnostic image.

The following is a summary of the steps to success when using AEC:

The performance of any automatic timer is dependent on the knowledge and skill of the technologist.

Positioning is crucial when using AEC devices.

Choose the detectors that are directly below the dominant area of interest.

The detectors cannot differentiate between secondary and scattered

radiation. Therefore, proper collimation is very important.

The back-up timer is a safety feature and should be set approximately 2-3x greater than the estimated mAs of the exposure.

Density changes should only be made by changing the density selector.

When using AEC, kVp changes should only be made to vary contrast.

To decrease the density of dark films due to long minimum response times, lower the mA.

If used properly, automatic timers can decrease repeat rates and increase department efficiency.

Bohr's Atomic Model of the atom

A basic explanation of these processes can be accomplished without using quantum

theory. A short reference to Bohr's model of the atom will suffice. According to this

model, an atom consists of a heavy nucleus and a number of electrons arranged on

well defined shells around this nucleus. With increasing distance from the nucleus,

these shells are designated with the letters K, L, M, N, O, P, Q, etc. All nuclei, except

that of regular hydrogen, contain besides the positively charged protons an almost

equal number of charge free neutrons. The number of protons in the nucleus

corresponds to the element number of the material. In an electrically neutral atom, the

number of protons and the number of electrons are equal. The closer the electrons are

to the nucleus, the tighter they are bound to the nucleus by its' electric field, or (in

other words) the more energy is needed to push them out of their place on the shell.

Boomerang Filter

Definition

A filter in the shape of a boomerang made of an impregnated silicone material with a

similar density to muscle tissue

The filter is typicaly placed between the patient and the cassette with the thicker part

overlying the less thick patient part, as illustrated in the picture with a shoulder

examination.

Other uses include.

Lower maxilla,

D1 to D1 in the AP position

Cephalometry

Bremsstrahlung Radiation

If an incoming free electron gets close to the nucleus of a target atom, the strong

electric field of the nucleus will attract the electron, thus changing direction and speed

of the electron. The electron looses energy which will be emitted as an X-ray photon.

The energy of this photon will depend on the degree of interaction between nucleus

and electron, i.e. the passing distance. Several subsequent interactions between one

and the same electron and different nuclei are possible. X-rays originating from this

process are called bremsstrahlung. Bemsstrahlung is a German word directly

describing the process: "Strahlung" means "radiation", and "Bremse" means "brake".

The process can create photons of

practically all energy values between

zero and the maximum determined by

the total kinetic energy of the

incoming electron. The chances for the

generation of a photon with a certain

energy by this process decreases with

increasing energy and reaches

practically zero for the very unlikely

event that an incoming electron looses

all its energy in one single interaction.

Consequently, the resulting radiation

contains photons of practically all

energy values between zero and the maximum. The distribution of the relative number

of photons with a certain energy, as a function of that energy, will decrease with

increasing energy and will reach zero at the maximum energy. This is equal to the

energy the electron picked up during the acceleration by the electric field between

cathode and anode. This energy is conveniently measured in electron volts: one

electron volt (eV) is the energy acquired by an electron traveling through a potential

difference of one volt. Therefore, accelerating electrons in an X-ray tube with a

voltage of x kV will yield electrons with an energy of x keV, and this will also be the

maximum energy an X-ray photon emitted by this tube can have.

Characteristic (HD) Curve D v LogE Curve for Film

Definition

Characteristic curve, a curve used to show the exposure properties of a film or a film

screen system. The characteristic curve, which was described in 1890 by Hurter and

Driffield, is a representation of how the exposure of the film is related to the

measurable signal, i.e. the blackening of the film, or film density. The characteristic

curve is different for different film types but has a general shape as shown in Fig.1.

The base and fog density is measured on an unexposed film. The shape of the

characteristic curve tells the user the contrast properties (slope of the linear part) and

the useful exposure range (length of the linear part). It also will indicate the speed of

the film (or film-screen system), which can be judged from the curve's position along

the horizontal axis. The speed class can also be found from the characteristic curve.

Average gradient,

the slope of the linear part of the characteristic curve of an X-ray film. This is

normally defined from the characteristic curve of an X-ray film using the density

points 0.25 and 2.0 (over the film fog) to calculate the average gradient as the slope:

(2.0 - 0.25) / (logE2.0 - log E0.25) where Ex is the exposure (or mAs) needed to

produce density x (over the fog level).

Fog,

film density caused by the development of silver halide grains that are not exposed to

light or X-rays in the exposure of the patient. There are many reasons why fog is

present in all X-ray images. Some of them are:

-- chemical fog

-- storage of the film in warm and humid locations

-- storage of film in locations with a high level of background radiation

-- contaminated developer solution

-- too high temperature and/or prolonged time in developer

(Kodak)

Characteristic Radiation

If the energy of the incoming electrons exceeds the binding energy of the electrons on

a certain shell of the target atoms, an additional process can happen: In a collision, the

incoming electron (1) can push the target electron (2) out of its place on the shell.

This event will leave an unstable atom behind. The gap on the shell will be filled

immediately by an electron (3) from an outer

shell or even from the conduction band of the

target material. This replacement electron

will thereby change its energy by a well

defined amount depending on the binding

energy levels of the electrons in the target

material, which are characteristic for that

material. The resulting X-rays (4) with very

distinct photon energy values are therefore

called characteristic radiation. As the binding

energy values for the outer shells are not high

enough for most elements to generate

photons of noticeable energy, usually only

characteristic radiation generated by

electrons jumping into the K-shell is

considered.

Chemical Automixer

The easiest way to deliver and mix x-ray chemicals is to utilize an automatic chemical

mixing station, or an automixer.

You use an automixer in much the same manner that you would use replenishment

tanks. Automixers are really dual mixers in that they have separate mixing systems:

one for developer and one for fixer. Both mixing systems work in an identical manner.

There are several different specific gravity-type automixers on the market, all of

which do basically the same thing in the same manner. Since the automixer can mix

both developer and fixer, they have two sides: a developer side and a fixer side.

Automixers have one common electrical connection, but everything is separate from

there. Automixers are designed to fit the bottles that you are using into �templates�

so that you can�t accidentally put developer parts on the fixer side and fixer parts on

the developer side. You will remove the cap and place the bottle upside-down in the

�template.� (The bottles are sealed with a foil seal which prevents the bottle from

leaking when turned upside down in addition to protecting the chemicals from

exposure to air.) When inserted into the template upside-down, a knife mechanism

uses the weight of the bottle to cut the foil seal and release the liquid in the bottle into

the mixer. When the chemical enters the water below from the bottle you just inserted

into the template, (remember that water has a specific gravity of 1.000 and the

concentrated solutions you are mixing are much heavier than that) the specific gravity

of the resulting solution increases to a pointwhere a specific gravity float-switch is

lighter than the solution and begins to float. When the floatswitch is boyant (floating)

an electrical connection is made in the switch which opens a solenoid valve allowing

water to flow into the mixture. As the water enters the mixture, the solution gets

lighter and lighter (closer to 1.0) until the float-switch no longer has the ability to float.

The float-switch sinks which breaks the electrical connection in the switch and returns

the solenoid valve to its resting state which is normally closed, shutting off flow of

water. The float-switch is the key to accurate mixing. Each float-switch is custom

made by the mixer manufacturer and most are made to be adjusted over a wide range

of gravities. Most x-ray chemicals are designed to be mixed to have a working

strength specific gravity of 1.075 and 1.085. In addition, at the proper specific gravity

most x-ray chemicals are designed to mix to exactly five gallons. So, by measuring

specific gravity with a calibrated specific gravity float, you will be mixing accurately

to five gallons if the formula was designed for five gallons. The above process is

extremely easy, and except for placing the bottles on the mixer, the process is all

automatic and can be very clean. In addition, automatic chemical mixers help to

reduce the odors often associated with x-ray chemicals which makes for a more

pleasant and potentially healthier working environment. There are some negative

features to using an automixer. Automixers are not always one hundred percent

accurate. In most cases you wont see extreme variations from mix to mix, but the

accuracy of automixers is dependant on incoming water pressure variability, and

incoming water temperature. Water pressure which varies widely from time to time

can affect the way the chemicals are mixed by affecting the motion of the specific

gravity floatswitch. Variable incoming water temperature will affect the specific

gravity of the mixture itself. As temperature increases, molecules expand causing the

solution to lose density or become lighter. As temperature decreases, molecules get

closertogether which causes the density to increase or get heavier. Water temperature

will not change rapidly from mix to mix, or even from day to day, but in more

northern climates, water temperature can vary from 35 degrees F. in the winter to 70

degrees F. or higher in the summer. In theory, such a larger jump in temperature from

one season to the next will cause the solutions to become heavier in the winter and

lighter in the summer and the mixer should be calibrated twice per year as a result.

Most mixers don�t get calibrated after installation because the variability in chemical

density as a result of incoming water temperature is not great enough to cause

objections to film readability or processing quality. However, in some cases, you may

find a heavier build-up of chemicals in the processor or even in the mixer during the

winter months as a result of more densely mixed chemicals. The biggest negative

feature to using an automixer is really more of a hurdle than a negative feature:

someone has to put the bottles on the automixer. X-ray departments often see this as

causing more work when they are already overworked. It does take time to place the

bottles on the automixer, but probably no more than 60 seconds. Once the x-ray

department realizes that the work involved in using an automixer really is no work at

all, they love to use them. Using an automixer means being able to store more

concentrated chemicals which decreases the risk of running out of chemicals. In

addition, because the chemicals are delivered in bottles sealed in boxes, delivery and

storage is much neater and more convenient. In addition, the chemicals are delivered

in the manner that the chemical manufacturer has specified and you can depend of

quality chemical manufacturers to provide consistency from case to case and from

batch to batch and from year to year. So, an automixer allows the x-ray department or

facility the freedom from most mistakes and the freedom to have cleaner more

pleasant working environment.

Collimator

Typical basic collimator specification

EXTERNAL COVER IN A.B.S. PLASTIC. Multilayer, square field X-ray collimator for stationary units.

The collimator enclosure is constructed of double steel, lead-lined walls for maximum X-ray

protection .

X-ray field size is limited by 6 pairs of shutters four of which are lead-lined. Two pairs of shutters are positioned near the focus, two near the entrance window and two near the exit window of the X-ray beam from the collimator. The shutters are controlled by two knobs located on the front panel. Besides the minimum in-built filtration, three variable filters may be added manually.

SPECIFICATIONS:

External adjustment of mirror angulation.

Additional filtration:

0,1mm copper in addition to the 1mm Al. Support 0,2 mm copper in addition to the 1mm Al support. 1mm aluminium in addition to the 1 mm Al support High luminosity provided by a quartz iodide lamp. Timer limiting projection lamp exposure time to 30 seconds, adjustable, thus extending lamp life and preventing overheating. 150 kVp radiation shielding. Minimum inherent filtration 2mm aluminium equivalent. (0,3mm and 1mm on request)

Continuous film coverage from 0 x 0 to 43 x 43 cm, 1% FFD, at 90cm FFD (SID).

Notes

When it comes to radiation protection, the collimater plays an important

role:

It is used to narrow the radiation field to a size needed for the examination at hand. For this it is equipped with sets of lead plates providing either a round or a square-shaped radiation field.

These collimating plates are either motorized or operated manually. In automatic mode, the image-receiver size is detected, and the collimating plates are operated accordingly

Diagrams of collimator with (1) no light beam (2) Light beam

Compound Anode

Compound anodes are used where there is a requirement for a high thermal capacity.

Compound anodes typically have a target area of rhenium tungsten alloy backed by

molybdenum and or graphite, the tungsten rhenium alloy resists surface damage and

"crazing" better than a pure tungsten surface, the molybdenum and or graphite and

molybdenum has a higher heat capacity than a purely tungsten anode.

The greater the volume of metal the higher the heat capacity, molybdenum has half

the weight of tungsten and graphite one tenth the weight so a greater mass can be used

with no increase in weight and its associated problems for rotation, construction etc.

c/o Phillips Medical

Compton Scattering

Compton Scattering, also known as incoherent scattering, occurs when the incident

x-ray photon ejects a electron from an atom and an x-ray photon of lower energy is

scattered from the atom. Relativistic energy and momentum are conserved in this

process below) and the scattered x-ray photon has less energy and therefore a longer

wavelength than the incident photon. Compton scattering is important for low atomic

number specimens.

The change in wavelength of the scattered photon is given by:

Theta is the scattering angle of the scattered photon. Note the fundamental constants

for the speed of light, Planck constant, and electron mass.

Contrast (Image)

Contrast is the difference in density between the darkest and lightest areas of the

subject image.

In the diagarm below contrast is the difference in the length of line b and line a.

Comparing contrast between films for the same subject

In the diagram below it is obvious the the difference between density a is greater than

a' there for between the two films Film A has greater contrast

Diagnostic Reference Levels Working Party

Statement

IPEM/NRPB/RCR/CoR/BIR Diagnostic Reference Levels Working Party

The Ionising Radiations (Medical Exposures) Regulations 2000 requires employers to

establish diagnostic reference levels (DRLs) for radiodiagnostic examinations. The

Institute of Physics and Engineering in Medicine (IPEM) along with the National

Radiological Protection Board (NRPB), the College of Radiographers (CoR), the

Royal College of Radiologists (RCR) and the British Institute of Radiology (BIR)

have established a Working Party to provide guidance on the implementation of DRLs

for diagnostic x-ray examinations. The membership of the Working Party is:

Mr A Workman (IPEM), Dr J Kotre (IPEM), Mr A Shaw (IPEM), Ms R Fong (IPEM),

Mr B Wall (NRPB), Dr R Bury (RCR), Mrs S Barlow (CoR), Dr D Sutton (BIR), Mr

J Williams (BIR) and Mr S Ebdon-Jackson (DoH observer)

The Working Party had its first meeting on 5 October and has issued the following

preliminary guidance.

Diagnostic Reference Levels should be seen as part of the overall framework for

protection of the patient along with the other requirements of IR(ME)R 2000 and of

Regulation 32 of IRR 99 .

Membership

Meetings

Publications

Projects

News

Reports

Links

How to join

Register

1. National DRLs

A Department of Health working party on DRLs, which included representation from

professional bodies and other organisations associated with medical exposures, met

on 13 January 2000. At this meeting it was agreed that for diagnostic x-ray

examinations the rounded third quartile values from the 1995 NRPB patient dose

review (NRPB-R289) [1] would be proposed as National DRLs. The values and

examination types are as follows:

Radiograph/Examination

National Diagnostic reference level

Examination Entrance surface dose

(mGy)

Skull AP/PA 4

Skull LAT 2

Chest PA 0.2

Chest LAT 0.7

Thoracic spine AP 5

Thoracic spine LAT 16

Lumbar spine AP 7

Lumbar spine LAT 20

Lumbar spine LSJ 35

Abdomen AP 7

Pelvis AP 5

IVU 25

Barium meal 17

Barium enema 35

It was further agreed that National DRLs would be reviewed at five-yearly intervals,

and that individual medical physics departments and hospitals carrying out

programmes of patient dosimetry in diagnostic radiology should be strongly

encouraged to contribute data to NRPB for national collation. The list of examinations

for which there are national DRLs will be extended when sufficient data on UK

practice has accrued.

2. Local DRLs

IR(ME)R requires the Employer to establish DRLs for radiodiagnostic examinations.

Employers should adopt a set of DRLs, having regard to national and European DRLs

where available. In the first instance, the examination types and DRL values adopted

can be those of the established national DRLs. Where a national or European DRL is

not available for an examination type, there is no requirement to set a DRL locally for

this examination. Any relevant local patient dosimetry data should be reviewed to

identify examinations where established local practice will support the adoption of a

DRL value lower than the equivalent national DRL. The local adoption of a DRL

which is higher than the respective national value will need to be justified. For

example, a case-mix which consistently requires examinations of greater duration and

complexity than the norm may justify a higher patient dose. However, a local DRL

higher than the equivalent national value cannot be justified solely on the grounds of

the use of poor equipment and/or techniques. A hospital or Trust Radiation Protection

Committee, Medical Exposures Committee or their equivalent would be a suitable

forum for ratifying locally adopted DRLs.

Local adoption of DRLs makes employers responsible for the level at which the

DRLs are set, in line with the concept of Clinical Governance. It is important to note

that this does not mean that individual Trusts must derive their own DRLs from their

own locally measured patient doses. Local measurements from one Trust may not

produce statistically valid DRLs.

The Department of Health working party meeting in January 2000 agreed that local

DRLs should be reviewed annually. Annual review of DRLs is intended to provide a

formal mechanism for revision of locally adopted DRL values which may follow

revised or new national DRLs, or additions to local patient dose data. Where

examination protocols have been changed, the effect on the locally adopted DRLs

should be considered, but it is not intended that this review requires annual patient

dose surveys.

3. Reviews triggered by DRLs being 'consistently exceeded'

Employers are required to undertake a review if a DRL is ‗consistently exceeded‘.

Because of the known wide variability in doses between individual patients for the

same type of examination, DRLs are defined as dose levels for typical examinations

for groups of standard-sized patients (or standard phantoms). Therefore comparing the

dose to an individual patient with a DRL has limited value. Rather, the distribution of

doses on a representative group of close to standard-sized patients (or on a standard

phantom) should be considered. For example, the mean value of this distribution can

be compared with the DRL to determine whether a DRL is being ‗consistently

exceeded‘.

Regulation 32 of IRR 99 requires that a suitable quality assurance programme be

provided for equipment used for medical exposures, which should include periodic

measurements of representative doses to samples of average size patients. The basis

of a suitable quality assurance programme is outlined in IPEM Report 77 [2]. This

incorporates national recommendations for patient dosimetry [3] which state that

measurements should be made at least every 3 years on each piece of equipment or

whenever changes are made to equipment or procedures that are likely to significantly

affect patient dose. The Working Party believes that these periodic patient dose

assessments required by IRR 99 can usefully be used to determine whether DRLs are

being consistently exceeded.

The average dose to a group of standard-sized patients measured by such surveys

should be compared to the respective locally adopted DRL. Where the DRL is

exceeded, the employer must instigate a review of local practice to establish reasons

and implement corrective action, where appropriate. It is expected that this ongoing

audit of compliance with DRLs can be achieved in most Trusts by the existing rolling

programme of 3-yearly patient dose measurements. This patient dose assessment

programme is distinct from the annual review of locally adopted DRLs discussed in

section 2.

The Working Party intends to consider the application of DRLs to other types of x-ray

examination, and to provide practical guidance on the required review processes.

References

Doses to Patients from Medical X-ray Examinations in the UK - 1995 Review. NRPB

R289, 1996

Recommended standards for the routine performance testing of diagnostic x-ray

imaging systems. IPEM Report 77, 1997

National protocol for patient dose measurements in diagnostic radiology. IPSM, CoR

and NRPB 1992

Digital Flat Panel (From the GE website)

The third in the series of the Revolution™ Digital Flat Panel Education will

cover the specifics of the GE Revolution™ Digital Flat Panel Detector.

Discover the principles, technology and design structure of the GEs Digital

Flat Panel in this issue.

GEs Digital Flat Panel Detector History

Result of extensive corporate R&D since

1985, the GE Revolution™ Digital Flat Panel

(DFP) detector replaces the film in

Mammography and Rad applications as well

as the analog image intensifier, with its

camera optics, pickup tube or CCD camera,

and analog-to-digital converter, in Cardio-

Vascular applications.

Using a common technology platform that requires only limited customization for

each application, GE pioneered the deployment of DFP detectors in Mammography

(1999), in Rad (1999) and in Cardiac (2000).

Characterized by a very high Detective Quantum Efficiency, the GE Revolution™

detector captures nearly all the information available at its entrance and transfers it

with almost no degradation to the observer. For all applications, the result is

outstanding image quality at reduced dose.

GEs Digital Flat Panel Technology

The principle of the flat-panel detector is illustrated in the drawing below.

The cesium iodide (CsI) scintillator absorbs x-ray photons, converting their energy

into light photons emission. This light is then channeled toward the amorphous silicon

photodiode array where it causes the charge of each photodiode to be depleted in

proportion to the light it receives. Each of these photodiodes is a picture element

(pixel); the spatial sampling of the image, which is the first step in image digitization,

is thus performed exactly where the image is formed, whereas it is realized almost at

the end of the chain in an Image Intensifier (see more in part 2 of the education series).

The electronic charge required to recharge each photodiode is then read by ultra-low-

noise proprietary electronics and converted into digital data that are then sent to a

real-time image processor. In the GE cardiac system, over 30 million pixels per

second are read out, processed, and displayed in real time.

GEs Digital Flat Panel Structure

The heart of the flat panel digital detector

consists of a two-dimensional array of

amorphous silicon photodiodes and thin-film

transistors (TFTs), all deposited on a single

substrate.

Utilizing thin film technology similar to that

used in the fabrication of integrated circuits,

layers of amorphous silicon and various

metals and insulators are deposited on a glass

substrate to form the photodiodes and TFTs

matrix, as well as the interconnections, and the contacts on the edges of the panel.

Principle of the GE Revolution™ Digital Flat Panel Detector.

Mono-substrate Amorphous Silicon panel coated with CsI

scintillator.

The CsI scintillator, which converts x-ray photons into visible light photons, is

deposited directly on top of the amorphous silicon structure.

Using a proprietary process, it is grown in very thin needles (5µm width) that channel

the light photons towards the photo-diode, like a fiber optics would do. This allows

one to increase the thickness of the CsI, and thus to stop and detect more X-rays,

without degrading spatial resolution because of wide-spreading light scatter as

observed in typical radiographic phosphor screens.

The photo-diode comprising each pixel is used as a bucket for electrons and each TFT

behaves as a switch to access the associated photo-diode. The TFT conductive state is

controlled through the voltage applied by scan electronics modules to matrix rows.

When a TFT is conductive, the charge of the corresponding photo-diode can be

measured through a matrix column by the readout electronics modules and converted

to a digital value by the analog to digital converter attached to each colomn.

The second step of image digitization after spatial sampling: pixel quantification, is

thus also performed next to image formation, and not at the end of a long

transformation chain like in an Image Intensifier-based system. (for more details on

Image Intensifier imaging chain, see part 2 in the Digital Flat Panel education series).

Electron microscope views of CsI needles that

constitute the scintillator layer.

Flat Panel and Imaging

Scan modules and readout modules are GE proprietary designs and use state-of-the-

art high density packaging technology to minimize sources of noise. Associated with

the optimized design of the amorphous silicon flat panel, the electronic noise

generated in the entire detection chain, from the photo-diode to the output of the

analog-to-digital converter, is equivalent to the signal generated by a single X-ray

photon. Thus, the read-out noise added by the panel is significantly less important

then the quantum noise in X-ray imaging. The image quality is therefore limited only

by the X-ray quantum noise, i.e. by the dose, and not by the detector performance.

This low noise performance, which is particularly important in fluoro where very low

dose is required. Combined with other advantages of the flat panel detector, such as

large dynamic, response stability over dose variations and time, response uniformity

over the entire image area, and absence of distortion, it provides a breakthrough in

image quality. All this adds not only to intrinsic image quality but also and opens new

opportunities for further image processing.

Processing Data with a Flat Panel

In cardio-vascular imaging, information is typically

associated with small objects such as arteries, stents,

guide wires, and catheters - objects that overlap each

other and large organs with different contrasts such

as lungs or diaphragm. Because the display has a

finite number of gray levels, representing the organs

at their acquired brightness levels may compromise

the representation of smaller objects of clinical

interest. At GE, we have developed state-of-the art computational methods to

represent the information in an intelligent manner, so that features of interest are

allocated optimal display values. This requires that the original image be captured

with high fidelity over its entire dynamic range.

As a result, the detector gives images a unique look and feel. This allows diagnostic

information to be presented with optimal utilization of display properties and human

visual perception. This technology also provides the ability to selectively enhance the

contrast of objects such as stents regardless of the anatomical background against

which they are acquired, providing better visibility of object details across the entire

image, regardless of the background anatomy.

Conclusion

The family of digital detectors manufactured by GE is based on a common technology

platform whose heart is a two-dimensional amorphous silicon array of photo-diodes

and thin-film transistors deposited on a single piece substrate and directly coated

with needle grown Cesium Iodide. The technology platform strategy forced the

design to be able to answer the most challenging needs of each application, such as

large field of view for chest Rad, high resolution for Mammography, real time and

low noise image acquisition for Cardiac. This strategy has several advantages:

fast introduction of the successive detectors customized for each application;

today, more than 1,100 systems are installed worldwide and give GE a unique

know-how in Digital detectors,

easy cross-fertilization between customized oanel formats and designs,

enables each customer to benefit from developments made to the panel for

other applications; this offers the ability to enjoy performances that exceed the

demands of today‘s practice and open the way to new breaking-through

applications.

The high performances of the amorphous silicon flat panel are complemented by the

proprietary electronics for detector control and readout. Associating one Analog-to-

Digital converter with each of the 1024 or 2048 pixels forming a single image row is a

good example of a design without compromise to minimize noise sources in all

conditions.

All that results in a final design which offers Image Quality performances as well as

simplicity, with a single large sensitive area requiring neither tile stitching with the

associated lost pixels, nor detector motion prohibiting fast acquisition, and thus a

reliability demonstrated by the most large and diverse installed base of Digital

Detectors

Diode / rectifier

The Diode Valve Invented in 1904 by John Flemming

A tungsten filament, similar to that found in an electric light bulb is heated by an

electric current, in a glass envelope containing a vacuum. This produces a cloud of

electrons around it, which are negatively charged. If a positively charged metal plate

is positioned near the filament, it will attract electrons from the cloud and a current

will flow in the circuit. If the plate is charged negatively the electrons will be repelled

and no current will flow. This is the basis of the valve rectifier. This type of valve is

called a diode, the filament is called the cathode and the plate is called the anode. The

diode valve was developed by Sir John Ambrose Fleming in 1904, while working for

the Marconi Company.

A diode with a filament as the cathode is called a directly heated valve, but most

modern valves are indirectly heated. An indirectly heated cathode consists of a nickel

tube that is coated with an oxide consisting of barium, strontium and calcium. Inside

the tube is the heater which consists of an insulated tungsten filament. This type of

cathode has two advantages. It produces more electrons and can be operated from an

AC or DC voltage.

Electronic symbols for the diode

The graph below shows that voltage current relationship is not a straight on off, in

reality the curved bit at the bottom is where more and more of the electrom cloud is

attracted across this is called the non saturated mode then when all the electons boiled

off are attracted as quickly as they are boiled off the graph is flat

Voltage current graph for a typical valve

The diode is used as a rectifier to convert an AC voltage to DC. In the circuit

below, the diode only conducts when the anode is positive with respect to the cathode,

and so only conducts on each positive half cycle of the AC input. The voltage at the

cathode consists of just the positive half cycles.

The capacitor charges to the peak of these half cycles to produce a smooth DC voltage.

DC output voltage shown in the graph, is not very smooth because the capacitor starts

to discharge between each positive peak. This is called the ripple voltage, and it can

be reduced by increasing the size of the smoothing capacitor

The diagram below shows how four diode rectifiers may be arranged in a "bridge" to

produce a direct current from an alternating current

The graph below shows the unsmoothed output waveform from a bridge rectifier

circuit

Dose Area Product Meter (DAP)

Dose-Area-Product (DAP) meters are large-area, transmission ionization chambers

and associated electronics. In use, the ionization chamber is placed perpendicular to

the beam central axis and in a location to completely intercept the entire area of the x-

ray beam. The DAP, in combination with information on x-ray field size can be used

to determine the average dose produced by the x-ray beam at any distance

downstream in the x-ray beam from the location of the ionization chamber. The use of

DAP is discussed further later.(2)

A recent modification of the ionization chamber design used in a DAP meter has

resulted in an instrument that measures both DAP and the dose delivered by the x-ray

beam. This design effectively combines data from a small ionization chamber that is

completely irradiated by the beam and independent of the collimator adjustments with

the conventional DAP meter.

Some fluoroscopic and radiographic systems have dose-area product (DAP) meters.

DAP meters measure the radiation dose to air, times the area of the x-ray field. The

relationship between DAP and exposure-area product (EAP) is essentially a single

conversion factor that relates dose to exposure. EAP is expressed in roentgen-cm2 (R-

cm2) and DAP is expressed ingray-cm2 (Gy-cm2).

How is DAP measured?

An ionization chamber larger than the area of the x-ray beam is placed just beyond the

xray collimators. The DAP ionization chamber must intercept the entire x-ray field for

an accurate reading, one proportional to the EAP. The reading from a DAP meter can

be changed by altering the x-ray technique factors (kVp, mA, or time), varying the

area of the field, or both. If the chamber area is larger than that of the collimators, as

the collimators are opened or closed the charge collected will also increase or

decrease in proportion to the area of the field. For example, a 5 x 5 cm x-ray field

with an entrance dose of 1 mGy will yield a 25 mGycm2 DAP value. If the field is

increased to 10 x 10 cm, with the same entrance dose of 1 mGy the DAP increases to

100 mGy-cm2, which is 4 times the DAP for the 5 x 5 cm field.

DAP meter Sensor DAP meter display with print out (Vertec)

Why DAP?

Dose-area product is relatively easy to measure. DAP meters have been around for

many years, and were actually used in the 1964 and 1970 U.S. X-ray Exposure

Studies. Advocates of DAP meters contend that the DAP is a better indicator of risk

than entrance dose alone, since DAP incorporates the entrance dose and field size.

DAP has been shown to correlate well with the total energy imparted to the patient,

which is related to the effective dose and therefore to overall cancer risk.

Are there problems with DAP?

There are several problems with the use of the DAP value. The configuration of the

DAP meter may introduce a bias to the DAP value. For example, if any material is

placed between the meter and patient, the patient will receive less than what is

implied by the displayed DAP value. For an undertable fluoroscopy system this can

be the tabletop and pad. Consequently, the use of DAP to estimate skin entrance

exposure or skin dose is complex and should only be attempted by a qualified medical

physicist. This is particularly true for fluoroscopic procedures where multiple beam

directions, source-skin distances, and field sizes may be used. DAP meters are

difficult to calibrate and maintain. Large changes in the DAP meter response can

occur over time, particularly if meters are adjusted for couch transmission factors.

Calibration should be done in the field after any changes that might alter the DAP and

at least annually.

2. Dose-Area-Product Problems in usage

Up to a decade ago, radiological patient safety concerns were focused on stochastic

risk. Monitoring and managing stochastic risk requires estimates of the effective dose

delivered to the patient. There is no need for real-time feedback. DAP is defined as

the integral of dose across the X-ray beam. Therefore DAP includes field non-

uniformity effects such as anode-heel-effect, and the use of semi-transparent beam-

equalizing shutters (lung shutter). DAP is easy to measure. The simplest method is to

place a transmission full-field ionization chamber in the beam between the final

collimators and the patient. DAP may also be obtained by calculation. Data is

accumulated during fluoroscopy, fluorography, and radiography. Assuming that the

incident beam is totally confined to the patient, the recorded value essentially

provides an upper limit on the X-ray energy absorbed by the patient (i.e. there is no

transmission or scatter). DAP‘s ability to estimate stochastic risk is degraded because

of the lack of dose distribution information within the patient. The best that one can

do is to assume an average weighting factor for all the tissues at risk. This may lead to

an over or under estimate of risk in certain cases. As an example, it does not account

for the differential risk of breast cancer from an AP or a PA projection. DAP rate and

cumulative DAP can easily be displayed in real-time. The primary utility of DAP rate

is in a teaching situation. Scattered dose rate at any place in the lab is more or less

proportional to DAP rate. The trainee can be shown that reducing DAP rate reduces

his or her personal exposure rate. The effect of different control options (e.g.,

collimation, zoom mode) on DAP rate can be demonstrated. Cumulative DAP does

not provide a direct indication of the possibility of skin injury. The same DAP is

observed with large fields and low skin doses as with small fields and high skin doses.

Exceeding skin tolerance is more likely in the latter case. However, reasonable

entrance field size estimates can be made for many procedures. These estimates are

dependent on factors such as equipment configuration, patient size, and operator

technique. Once known, the nominal field size can be used to obtain an estimate of

skin dose. Rules-of-thumb can be established to make this conversion for typical

procedures. DAP provides no information regarding the spatial distribution of the

entrance beam around the patient‘s skin. It produces an overestimate of the possibility

of exceeding the deterministic threshold when there is significant beam movement

during the procedure.

Summary:

DAP meters are valuable quality control tools for monitoring changes in equipment

and procedures. DAP does not represent radiation dose per se, and use of a DAP

meter to determine patient dose should only be attempted by a qualified medical

physicist. DAP meters need to be recalibrated on a regular basis—at least annually—to

maintain adequate accuracy.

Dose Definitions

Absorbed dose: The energy imparted per unit mass by ionizing radiation to matter at a

specified point. The SI unit of absorbed dose is the joule per kilogram. The special

name for this unit is the Gray (Gy).

Air kerma: The energy released per unit mass of a small volume of air when it is

irradiated by an x-ray beam. For diagnostic x-rays, air kerma is the same as the

absorbed dose delivered to the volume of air in the absence of scatter. Air kerma is

measured in Gy.

Biologic variation: With respect to radiation, the differences among individuals in the

threshold dose required to produce a deterministic effect, or the differences in degree

of effect produced by a given dose. Biologic variation may be idiopathic

or due to

underlying disease. Different areas and types of skin also differ in radiosensitivity.

C-arm fluoroscopic system: A fluoroscopic system consisting of a mechanically

coupled x-ray tube and image receptor. Such systems typically have two rotational

degrees of freedom (left-right and cranial-caudal). Most such systems have an

identifiable center of rotation called the isocenter. An object placed at

the isocenter

remains centered in the beam as the C-arm is rotated.

Cumulative dose (CD): The air kerma accumulated at a specific point in space relative

to the fluoroscopic gantry (the interventional reference point) during a procedure. CD

does not include tissue backscatter and is measured in Gy. CD is sometimes referred

to as cumulative air kerma.

Deterministic effect: A radiation effect characterized by a threshold dose. The effect is

not observed unless the threshold dose is exceeded. (The threshold dose is subject to

biologic variation.) Once the threshold dose is exceeded in an individual,

the severity

of injury increases with increasing dose. Examples of deterministic effects include

skin injury, hair loss, and cataracts.

Dose: As used in this document, "dose" is the same as the absorbed dose unless

specified as "equivalent dose" or "effective dose."

Dose–area–product (DAP): The integral of air kerma (absorbed dose to air) across the

entire x-ray beam emitted from the x-ray tube. DAP is a surrogate measurement for

the entire amount of energy delivered to the patient by the beam.

DAP is measured in

Gy·cm2.

Effective dose: The sum, over specified tissues, of the products of the equivalent dose

in a tissue and the tissue weighting factor for that tissue. Effective dose is measured in

Sieverts (Sv). Stochastic risk factors are usually stated relative to

effective dose.

Equivalent dose: A quantity used for radiation protection purposes that takes into

account the different probability of effects that occur with the same absorbed dose

delivered by radiations with different radiation weighting factors. Effective dose is

measured in Sv.

Fluorographic image: A single recorded image obtained using an image intensifier or

flat digital panel as the image receptor. A digital angiographic "run" consists of a

series of fluorographic images.

Fluoroscopy time: The total time that fluoroscopy is used during an imaging or

interventional procedure.

Interventional reference point (IRP): For C-arm–type fluoroscopic systems with an

isocenter, the IRP is located along the central ray of the x-ray beam at a distance of 15

cm from the isocenter in the direction of the focal spot. The IRP is defined by

International Electrotechnical Commission (IEC) standard 60601-2-43

Isocenter: For C-arm–type fluoroscopic systems, the point in space between the focal

spot and the image receptor through which the central ray of the x-ray beam passes,

regardless of beam orientation.

Kerma: Kinetic Energy Released in Matter; the amount of energy transferred from the

x-ray beam to charged particles per unit mass in the medium of interest. For

diagnostic x-rays, this is equivalent to absorbed dose in the specified medium (eg,

air,

soft tissue, bone). Kerma is measured in Gy.

Peak skin dose (PSD): The highest dose at any portion of a patient‘s skin during a

procedure.

Stochastic effect: A radiation effect whose probability of occurrence increases with

increasing dose, but whose severity is independent of total dose. Radiation-induced

cancer is an example.

Threshold dose: The minimum radiation dose at which a specified deterministic effect

can occur. Threshold doses differ among individuals as a result of biologic variation.

The threshold dose for skin injury also differs in different anatomic sites

on the same

individual.

Although practicing physicians should strive to achieve perfect compliance, in

practice, all physicians will fall short of this ideal to a variable extent. Indicator

thresholds may be used to assess the efficacy of ongoing quality-improvement

programs. For the purposes of these guidelines, a threshold is a specific

level of an

indicator that should prompt a review. When compliance rates fall below a minimum

threshold, a review should be performed to determine causes and implement changes

if necessary. If recording patient radiation dose data is one measure of the quality of

radiation dose management, compliance rates lower than the defined threshold should

trigger a review of policies and procedures within the department to determine the

causes and implement changes to improve quality. Thresholds may vary from those

listed here; for example, patient referral patterns and selection factors

may dictate a

different threshold value for a particular indicator at a particular institution. Because

institutions and interventional fluoroscopic units vary widely in their ability to

measure various metrics of patient dose, radiation dose data may be recorded

with use

of one or more of four different dose metrics: fluoroscopy time/number of

fluorographic images, DAP, CD, and PSD. Therefore, setting universal thresholds is

very difficult and each department is urged to alter the thresholds as needed to higher

or lower values to meet its own quality-improvement program needs.

Exposure Factors

Kv / Kvp Potential difference between film and anode

The energy (you can consider this the penetrating power) of the x-ray beam is

controlled by the voltage adjustment. This control usually is labelled in keV (thousand

electron volts) and sometimes the level is referred to as kVp (kilovoltage potential).

Do not be confused by the different terminology, just remember there is a control by

which the difference in potential between the cathode and anode can be controlled.

The higher the voltage setting, the more energetic will be the beam of x-ray. A more

penetrating beam will result in a lower contrast radiograph than one made with an x-

ray beam having less penetrating power. It is probably obvious that the more energetic

the beam, the less effect different levels of tissue density will have in attenuating that

beam.

The generator waveform if is not constant potential (medium frequency etc) will

affect the effective Kv.

mA Tube Current

The second control of the output of the x-ray tube is called the mA (milliamperage)

control. This control determines how much current is allowed to flow through the

filament which is the cathode side of the tube. If more current (and therefore more

heating) is allowed to pass through the filament, more electrons will be available in

the "space charge" for acceleration to the target and this will result in a greater flux of

photons when the high voltage circuit is energized. The effect of the mA circuit is

quite linear. If you want to double the number of "x" photons produced by the tube,

you can do that by simply doubling the mA. Changing the number of photons

produced will affect the blackness of the film but will not affect the film contrast.

S Time

The third control of the x-ray tube which is used for medical imaging is the exposure

timer. This is usually denoted as an "S" (exposure time in seconds) and is combined

with the mA control. The combined function is usually referred to as mAs or

milliampere seconds so, if you wanted to give an exposure using 10 milliampere

seconds you could use a 10 mA current with a 1.0 second exposure or a 20 mA

current for a 0.5 second exposure or any combination of the two which would result in

the number 10. Both of these factors and their combination affect the film in a linear

way. That is, if you want to double film blackness you could just double the mAs.

The X-Ray beam

The x-ray beam has two main properties you need to understand.

1) Beam QUALITY is the ability of the beam to penetrate an object, its all about the

penetrating power of the x-ray photons, this is controlled by the KV control.

2) Beam INTENSITY this is the number of x-ray photons in the beam and is

principally controlled by the mAS

But note as you increase the KV not only does the QUALITY harden (more

penetrating) but you do actually get more photons so INTENSITY increases too.

Putting it all together the exposure Any radiographic subject has a minimum Kv required for the x-ray photons penetrate

the most dense part of the subject, the most radiographicaly dense part of the subject

will depend upon what the part is chemicaly composed of (Atomic number) and its

thickness (remember linear attenuation coefficients and HVL!?)

The thicker the subject the more absorption of x-rays so the thicker the part the more

mAS you require.

In theory the more Kv you use the less the contrast of the image will have

However in practice film screen / processing conditions affect contrast much more

In practice it is not as simple as this as scatter is produced which is not image forming

but adds density to the film and needs to be controlled, if you remember all those

complex diagrams about interactions of x-rays with matter you will realise the amount

and direction of scatter depends on the Kv and the material absorbing the x-rays.

Image 1 The Kv is too low the femoral

condyle is under pentrated you cannot see

the bone trabecualr patterns. the contrast is

too high to demonstrate all the soft tissues.

Image 2 Much better the all the subject is

penetrated and all the soft tissues are

visible

A well exposed abdomen image

demonstrating all the soft tissue structures.

A good chest image the mediastinum is

pentrated the image is exposed well

demonstrating the bones and soft tissues.

Under penetrated

OK

Under penetrated

Too much mAS

Too Little mAS

A few myths

Changing the Kv by 2 or 3 makes almost no perceptable image change!

Adding 10 Kv does not double the image density

Exposure factors are an exact science !

(the image you produce must satisfy the radiologist who interprets the image - not all

radiologists like the same penetration / density / contrast for the same body part)

Image Contrast Here, we need to spend a little more time discussing the issue of radiographic contrast.

This is an important concept because image contrast plays a critical part in the

interpreter's ability to detect abnormalities which are only slightly different from the

density of the surrounding material. It is not possible to say what is the optimal

contrast (or the optimal radiographic technique) for all situations. Different body parts

have different inherent tissue contrast. This can be illustrated by using the extreme

examples of the chest and the breast. In the chest, there is good inherent tissue

contrast with densities ranging all the way from bone at the high end to air at the low

end. On the other hand, the breast is inherently very low in tissue contrast only

containing structures which are water density (glandular material or tumor) or fat

density. For the moment, we will disregard small calcifications which are really not

normal structures. Because of this difference in inherent tissue contrast, we would be

likely to use a very low contrast radiographic technique for the chest because we have

good tissue contrast. Conversely we would be likely to use a very high contrast

technique for the breast because the breast has minimal, inherent tissue contrast.

Remember, image contrast is controlled by the energy of the "x" photon beam.

Therefore, high kV techniques result in low contrast images (the assumption is always

made that the image will have approximately the same average film density so if kV

is increased, there must be a compensation in mAs to keep film density constant). To

increase image contrast in situations where there is low tissue contrast, a low kV, high

mAs technique should be used. This is obvious for mammography but you should also

remember this possibility for other special situations such as looking for low-density

foreign bodies embedded in soft tissue. To improve film contrast for mammograms

we would need to use a very low energy x-ray beam. Mammograms are frequently

done with beams in the 25 keV range. For the chest x-ray, we would like to use a low

contrast technique which requires a relatively high-energy beam. Chest x-rays are

frequently done with beam energies above 100 keV. You should understand that for

similar film densities, the high KV technique usually results in lower patient radiation

exposure. Think about this long enough to clearly understand why less radiation is

absorbed in the patient when a high-energy beam is used.

Grids

One of the problems in getting a sharply defined image in clinical radiology is the

presence of scattered or secondary radiation. These photons are created in the body of

the patient or closely surrounding objects by the interaction of that material and the

primary "x" photons coming from the x-ray tube. Several possible interactions occur

in the diagnostic energy range. At relatively low energies, the photoelectric effect is

probable. The photoelectric effect is actually the desirable, photon/tissue interaction

because there is complete absorption of the photon with no production of a secondary

photon. The more common tissue interaction at the photon energies used for the

majority of clinical procedures is called the Compton effect or coherent scattering. In

this interaction, a secondary photon is produced at the site of interaction. The

secondary photon will always have lower energy than the primary photon and will be

going in an altered direction. These secondary photons, if allowed to reach the film,

will actually produce erroneous information by recording gray tone variation (and

therefore indicating relative tissue densities) at some distance from the site at which

the photon/tissue interaction actually occurred. The net result of allowing a significant

number of secondary photons to reach the film is a reduction in image sharpness.

There will always be a loss of spatial resolution.

Several methods have been devised to reduce the problem of scattered radiation. The

simplest and most direct is to simply limit the field of exposure. If a small image area

is adequate to make the clinical diagnosis, the image area should be "coned down" to

that small size. For instance, if you want to image the gallbladder, you will get a much

sharper picture if you bring the shutters down to include an area only the size of the

gallbladder instead of including the entire upper abdomen on the image. Just

remember that the smaller the area of the x-ray beam the fewer scattered photons you

will produce.

In the typical clinical imaging situation, the most common method of reducing scatter

is to use a radiographic grid. The grid looks like a flat metallic plate the size of the x-

ray film if you look at it directly. However, it is more complicated than that. It

actually is composed of alternating radiopaque (lead) and radiolucent (aluminum)

strips. These are arranged on edge, sort of like looking at the strips of a venetian blind

which is arranged to let light come between the strips. The edge of these strips is

turned towards the source of x-rays and in the most commonly used grid, the focused

grid, the anglulation of the strips is arranged to match the divergence of the x-ray

beam.

This arrangement of the radiographic grid will give the highest probability for primary

"x" photons passing between the lead grid strips and reaching the film, while the off-

focus or secondary photons are likely to interact in the lead strips and never reach the

film.

The use of this radiographic grid will greatly improve image sharpness when a

relatively thick body part is being imaged. Unfortunately, there is always a trade off.

Since the grid does stop some of the photons which would contribute to film

blackening, if you just add a radiographic grid without changing the tube settings, the

film will be greatly underexposed. If you decide to use a grid, you will have to

increase the number of photons produced by the x-ray tube in order to get the correct

film exposure. This will result in giving the patient increased radiation exposure.

Remember, the position of the grid is between the patient and the film.

The third method of reducing scatter or at least reducing the probability that scattered

photons will reach the film is to use an air gap. This is infrequently used in clinical

radiography but can still, sometimes be used to an advantage particularly when

magnification of the image might be helpful. Ordinarily we would have the film

positioned as close to the patient's body as possible for the radiography of any body

part. With an air gap technique, the film is moved several inches away from the

patient's body. That separation, (because secondary photons are likely to be lower

energy and moving at a greater angle than primary photons) will result in a decreased

probability of the secondary photon hitting the film. From the diagram below, you

will be able to understand that creating the air gap will also result in magnifying the

radiographic image. Remember the x-ray beam is produced from almost a point

source and it diverges as it goes towards the patient.

Extra focal Radiation

Extra focal radiation is radiation produced at the amode which is not from the area

represented by the focal spot.

Ferlic Filter

· An external beam filtration device as specified below

Improves image quality by filtering.

· Even density of x-ray signal and image from C1 to T1.

· Even density on lateral hips from acetabulum to the distal femur.

· Perfect compliment to Lateral Hip Support .

· Reduce or eliminate Swimmers view retakes.

· Reduce overall patient exposure and medical staff dose.

· Increase patient flow rates (not waiting for medical staff to apply traction etc.).

· Reduce costs and eliminate retakes.

· With conventional film, works best with ―L‖ films

e.g. Kodak T-mat L, Fuji HRL etc., etc..

· Also works with Computerised Radiography (CR).

Film badge – Holder

The photographic film dosemeter is designed to measure doses from X, beta and

gamma radiations in terms of the radiation quantities specified by the Health and

Safety Executive (HSE). The film badge service is approved by the HSE under

Regulation 35 of the Ionising Radiations Regulations 1999.

The dosemeter consists of a photographic film (manufactured by Kodak) contained in

light tight wrapper. To cover the required dose range the film incorporates two

emulsions, of different sensitivities. It is uniquely identified by means of a number

which is stamped onto the film and wrapper. The dosemeter also bears the wearer's

name or a serial number, the establishment code number, the expiry date and an

optional personal identifier for each employee, e.g. department name or a works

number. We also provide 'wear and care' cards for each member of staff. These are

designed to help users understand more about how and why they should wear the

dosemeter. The wearer places the wrapped film in a plastic holder, which is supplied

by the NRPB on permanent loan. The holder contains a number of metallic and plastic

filters which are necessary to ensure that the dosemeters provide an adequate

measurement over a suitably wide radiation energy range.

When developed the film darkens in proportion to the amount of radiation energy

received. From the differing amounts of filtration we can gain information on the

energy of radiation causing the dose. Radioactive contamination of the film can be

readily identified.

Film dosemeter technical specification

Detection gamma rays x-rays beta particples

Dose range

measured 0.1 mSv to 10 Sv 0.1 mSv to 400 mSv 0.1 mSv to 10 Sv

Energy

range

detected

10 keV to 7 MeV for

Hp (0.07)

20 keV to 7 MeV for

Hp (10)

10 keV to 7 MeV for

Hp (0.07)

20 keV to 7 MeV for

Hp (10)

700 keV to 3.5 MeV

(Emax) for Hp (0.07)

Periods of

use

2, 4, 8, 13 weekly

(calendar issue

periods are also

available)

2, 4, 8, 13 weekly

(calendar issue

periods are also

available)

2, 4, 8, 13 weekly

(calendar issue

periods are also

available)

Special features of the film dosemeter

Energy discriminating dosemeter

Through the use of several filters, the dosemeter is able to provide information on the

type and energy of the incident radiation.

Contamination

Radioactive contamination of the film can be readily identified.

Physical record

The film forms a physical record of the dose received by the wearer. The processed

film is stored by the NRPB for at least five years and may be accessed by the

customer.

Film Badge Holder - Personnel radiation Monitor Badge

The film holder is constructed of impact resistant plastic and features a snap-tight

hinged door that allows for easy replacement of film packets.

Whole body or area badges come equipped with a sturdy metal clip.

Film Packet The film is wrapped in a black protective paper and then sealed in a vinyl covering

that shields the sensitive material from light induced exposure. Tearing or puncturing

the covering will expose the film, therefore, destroying the ability to interpret the

processed film

Image 1 Film Badge Holders

The film badge's multi-filter system is designed so that radiation will reach one

quadrant of the exposed film after penetrating three different filter areas (plastic,

cadmium and copper) and passing

through an open window. A cadmium (Cd) filter absorbs particles with energies less

than 2 MeV and photons with energies less than 150 keV. Exposure to photons with

energies more than 150 keV is determined by comparing film response in the

cadmium filter area with the equivalent response on a calibration curve developed

with a Cesium-137 source. The film area under the copper (Cu) and plastic filters are

used to determine radiation exposures from photons of energy levels less than 150keV.

Exposure to beta particles is determined from the film response in the open window

area (after correcting for response from other radiations, as measured under the Cd,

Cu, and plastic filters) using the appropriate calibration curves.

Image 2 Diagram of a Film Badge Holder and Film

All calculations are performed on the dosimetry service's state-of-the-art computer

systems using data from calibration curves and related film response measurements

for each

film processed. Uniformity is important in film irradiation. Since all calibration

measurements for the film are made in the designated filter areas, film packets must

be exposed inside the film holder. The dosimetry service will not report results for

film exposed outside the holder.

The film holder is constructed of impact resistant plastic and features a snap-tight

hinged door that allows for easy replacement of film packets. Whole body or area

badges come

equipped with a sturdy metal clip.

Film Kodak Type 2 Personal Monitoring Film packets. Type 2 film consists of a single film

base with a

fast (sensitive) emulsion on one side and a slow (insensitive) emulsion on the other

side. Therefore, a single film in a convenient-to-use packet is capable of monitoring

exposures from a vast array of radiation hazards.

Film-screen Speeds

The sensitivity of a film-screen combination depends on the film, the screen, the film

processing, and the beam quality, i.e. the spectrum of the X-rays exposing the film

screen combination. This explains immediately, why the sensitometry of a film-screen

combination with X-rays is a lot more complex than the sensitometry of a film with

light, and therefore is hardly ever done outside the manufacturer's laboratory:

1. The film-screen combination has to be exposed with a standardized spectrum. This

requires the use of a specified high voltage value, a specified high voltage waveform

(usually DC), a specified target composition, a specified filtration, all resulting in a

specified half-value layer.

2. While the film-screen combination has to be exposed with different dose values,

the operating parameters of the X-ray source (tube voltage, tube current, and exposure

time) must not be changed, as this is the only way to avoid measurement errors due to

spectral changes and due to the reciprocity law failure. Therefore, the dose can only

be varied by changing the distance between source and film-screen combination.

3. The film has to be processed under standardized conditions.

The speed of a film-screen combination is stated as the inverse of the dose (in Gy)

needed to obtain a film density of one above base plus fog, multiplied by 1000 Gy:

1000 Gy

SPEED = -------------------------

Dose for D = 1+Base+Fog

The speed is the quotient of two dose values, it does not have a dimension or unit

name attached to it. As the speed is inverse proportional to the dose requirements of a

film-screen combination, twice the speed is equivalent to half the dose and vice versa.

With this definition, the standard or universal film-screen combinations with calcium

tungstate phosphor used to have a speed of 100. With the modern rare-earth systems,

the speed of the standard screen is usually 200, i.e. the film-screen combination for

universal application requires 5 Gy (approximately 0.5 mR) for a film density of

one plus base plus fog. The speed values of the high resolution ("detail" or "fine")

resp. the high sensitivity ("high speed") film-screen combinations of one and the same

product line differ from the speed (and thus, dose requirement) of the standard

combination by a factor of two in either direction. Thus, a rare-earth "detail" film-

screen combination has a speed of 100, and a rare-earth "high speed" film-screen

combination has a speed of 400. These are typical values, but for special applications

screens with lower and higher speeds are available

Filters / Filtration

Filtration,

removal of parts of the X ray spectrum using absorbing materials in the X-ray

beam. The X-ray spectrum reaching the patient is filtered by attenuating

material in its path. Filtering of the beam is used in order to modify the

spectral or spatial distribution of X-rays, or both. Filtration is in principal

divided in two parts: inherent filtration and added filtration. Among those

filters added are compensation and equalization filters. Examples of

compensation filters are variations of the wedge filter, which is used to

compensate for the otherwise uneven X-ray fluence generated by objects with

a wide thickness variation, such as hands and feet. Equalization filters follow a

similar principle and are sometimes used to compensate for more irregular

absorption variations in the object, such as the mediastinum in a chest frontal

image.

Inherent filtration, the filtration of an X-ray beam by any parts of the X ray tube or tube shield

through which the beam must pass. The parts include the glass envelope of the

X-ray tube, the oil cooling the tube and the exit window in the tube housing.

The inherent filtration corresponds to approximately 0.5–1 mm of aluminium.

The total filtration of the X-ray beam before it reaches the patient consists of

the inherent filtration plus the added filtration.

Added filtration,

commonly metallic filters inserted into the X-ray beam. The inherent filtration

normally consists of the filtration of the X-ray beam from the glass envelope

of the X-ray tube, the oil cooling the tube and the exit window in the tube

housing. In excess of this, added filtration is almost always considered needed.

This filtration is for normal X-ray purposes commonly made of aluminium or

copper. The purpose of inserting such extra filtration into the X-ray beam

serves the following purposes:

· To remove the low-energy photons that never would have been able to reach

the film and produce an image. These photons would, if present, only increase

the radiation dose given to the patient.

· To remove those low-energy photons that otherwise would have reached the

film but would have given rise to too high contrast in the image. The classical

example of this is in chest imaging, where the contrast from ribs and shoulder

blades must be reduced.

In other cases, extra filtration can fulfil other purposes:

· In mammography, where a molybdenum anode is used, the added filtration

normally is made of the same material (molybdenum Mo ). It is a fact that a

material is particularly translucent to its own characteristic radiation, therefore

giving an X-ray spectrum with as much (monoenergetic) characteristic X-rays

from Mo as possible and filtering more selectively on both the high- and low-

energy side of the Mo characteristic X-rays.

· For very special purposes, special filters can be used that will create a shape

of the X-ray spectrum that — to some extent — will ―match‖ the absorption

characteristics of the X ray contrast medium, thereby selectively increasing

their contrast properties.

Coppper & Aluminium Filtration

All the radiation absorbed inside the body, without having a chance of

penetration and forming an image, is harmfull radiation only!

In order to make the radiation "less harmfull", filters are used. The soft

radiation is absorbed inside the filter while the hard radiation passes only

slightly effected.

As seen in the left graph, Aluminum attenuates the very soft radiation

drastically. The radiation spectrum shown is the result of 100kV tube voltage

in combination with a filtration equivalent to 2.5mm Aluminum.

According to international regulations, this is the minimum amount of

filtration and must be guaranteed by the tube assembly.

Additional filtration with copper can be employed to make the radiation

"safer". Notice the shift of the peak intensity to higher keV by absorbing the

lower energies. So, the radiation quality is hardened-up by increasing the

amount of filtration

Generators

X-Ray generators provide the tube current at the required voltage for x-ray production.

In a "perfect" case this would be a constant voltage, however transformers require

alternating voltages to work so some means of producing a constant voltage across the

x-ray tube from the rising and falling voltage produced by the high tension

transformer is required.

(Siemens)

Diagram showing typical waveforms from x-ray generator output circuits

Ripple The deviation of the voltage waveform across the x-ray tube from constant voltage is

named ripple

.the variation in the high-voltage expressed as the percentage of the maximum high-

voltage across the X ray tube during X-ray production:

Ripple factor (%) = 100 x (Vmax - Vmin)/Vmax

The ripple causes corresponding but relatively higher variations in the X-ray output. It

is an unwanted phenomenon in the X-ray production due to the lengthening of the

exposure time and the reduction in the average kV. The ripple is theoretically 100%

for the old-fashioned single phase X ray generator (in practice, it is less, however, due

to the smoothening effect of the high-voltage cable capacitance). The three-phase X-

ray generator have ripple factors in the range of about 3–25% (3-phase 6-pulse

generator: 13–25%, 3-phase 12-pulse generator: 3–10%). In the medium frequency

generator, the ripple factor decreases with increasing kVp. In this type of generator

(also named high frequency or inverter generator), the kV is controlled by adjusting

the frequency of the current prior to high-voltage transformation. Ripple is usually in

the range of 4–15%. There is practically no ripple in the constant potential X ray

generator

Constant potential x-ray generator,

An X ray generator providing a nearly flat high voltage waveform for the X ray tube.

The term may refer to

1) any generator providing high voltage with a ripple factor less than a certain limit,

e.g. 5%, or

2)* a special generator type briefly mentioned below.

1) A voltage ripple limit of 5% would include the 3-phase 12-pulse generator and the

medium frequency generator.

2)*The so-called constant potential X-ray generator is a very large and expensive

generator that provides the highest average X-ray energy of any X-ray generator type.

It is now used only for the most demanding applications. This generator uses a three-

phase line voltage coupled directly to the primary windings of the high-voltage

transformer, i.e. without an intermediate autotransformer. Regulation of the kilovolt

peak kVp and exposure time is done on the secondary (high voltage) side of the

transformer by means of high voltage electron tubes (triodes or tetrodes). The high

voltage supplied to the X-ray tube has a nearly flat waveform with a ripple less than

2%.

(http://www.amershamhealth.com/medcyclopaedia)

-------------------------------------------------------------------------------------------------------

-1) One pulse self rectified

In the simplest case the tube acts as a rectifier and the - a self rectified circuit

(Stylised graph of voltage v Time)

-------------------------------------------------------------------------------------------------------

-2) One pulse half wave rectified

In order to prevent the anode producing electrons as it becomes hotter and the

electrons flowing backwards and striking the filament a single rectifier can be placed

in series with the x-ray tube to ensure current flows only from filament to anode. -

(Stylised graph of voltage v Time)

Advantages and Disadvantages

Inefficient use of power no x-rays produced in negative half cycle

Possibility of reverse conduction - low power output - unless a rectifier is used

Minimum exposure time 0.02 S to include one whole AC cycle

However the unit can be made relatively small and cheaply for situations requiring

limited output.

-------------------------------------------------------------------------------------------------------

-3) Two pulse full wave rectified

(http://www.amershamhealth.com/medcyclopaedia)

A bridge rectifier circuit (see rectification) inverts the negative half cycles and double

the number of positive cycles are produced per unit time compared with a single

rectifier.

- the ripple is said to be 100%.

http://www.eleinmec.com/

Advantages and Disadvantages

The principal disadvantages are the inefficiency of radiation production due to the

pulsating waveform no providing enough voltage to produce x-rays for a portion of

the time, and the inability to select short exposure times.

-------------------------------------------------------------------------------------------------------

-4) Constant potential x-ray generator

The constant potential generator circuit has two capacitors across the output from the

rectifiers to smooth the pulsating waveform.

To further condition the waveform, there is a triode valve in series with each lead

and these control the output via a grid connection which has control signals fed to it

from a high resistance across the output of the triodes. Once the tube voltage has been

monitored via the high resistance potential divider it is constantly corrected by the

control unit attached to the triode valve control grids.

The triode valves also from part of the timer switching and can operate at

microsecond intervals with good accuracy. The high voltage supplied to the X-ray

tube has a nearly flat waveform with a ripple less than 2%.

Block Diagram of Constant potential x-ray generator circuit

Output waveform from the rectifiers and the smoothed output across the x-ray

tube

http://www.eleinmec.com/

Advantages and Disadvantages

High x-ray output per mAS

Smaller range of x-ray energies

Very small exposure times possible

However these generators are very expensive and tend to be large and have the

possibility of more to go wrong.

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-5) Three phase 6 Pulse Generator

Commercial electric power, the line voltage, is usually produced and delivered as

three phase alternating current. The period of each single phase may be 50 or 60 Hz.

The period of a 50 Hz AC has a duration of 1/50 s, or 20 ms. The three phase X ray

generator transforms and rectifies this AC into a high-voltage direct current (DC) with

either six or twelve forward pulses per 20 ms period. As compared to the 100% ripple

factor of single-phase generators, three-phase generators dramatically reduces voltage

ripple (13–25% for 3-phase 6-pulse, 3–10% for 3-phase 12-pulse).

X-ray production is therefore much more efficient. The so-called constant potential X

ray generator produces a voltage ripple less than 2% (hence the name), and it

produces the highest average X-ray energy of any X-ray generator type, with

exposure times less than 1 ms. This kind of generator is, however, very bulky, with

high costs and inefficient power consumption. The preferred modern generator today,

is therefore the almost equally efficient, much smaller and less costly medium-

frequency generator (also known as high frequency and inverter generator).

An X ray generator using a 3-phase alternating current (AC) line source, i.e. three

wires, each with a single phase AC that is one third cycle (120°) out of phase with the

other two (Fig.1). The three-phase transformer used in this generator has three sets of

primary windings and three sets of secondary windings, i.e. in effect three separate

high-voltage interconnected transformers. The three primary and secondary windings

are connected either in a wye1 configuration or a delta

2 configuration. In the three-

phase six-pulse generator, rectifiers in the high-voltage circuit produce two pulses for

each line, resulting in a total of six pulses.

Waveforms from the various circuit in a 3 phase unit parts

Star and delta

windings

(http://www.amershamhealth.com/medcyclopaedia)

Wye or Star configuration, 1

a star-shaped configuration or interconnection of the three windings in the primary or

secondary of a transformer in a three phase X ray generator.

Delta configuration, 2

one possible configuration of the windings in the primary or secondary side of a three-

phase transformer. The windings in this transformer can be arranged as a ――

Combinations of these configurations in the primary and secondary windings of a

transformer will give rise to a phase shift of 30. Using one delta and one wye

configuration as secondary windings and (usually) a delta configuration as primary

winding will therefore give twelve pulses per period of mains AC voltage.

Advantages and Disadvantages

High x-ray output per mAS

Smaller range of x-ray energies

However these generators are expensive and tend to be large and have the possibility

of more to go wrong.

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-

Three-phase twelve-pulse generator

In the three-phase twelve-pulse generator, a different configuration of transformers,

one of each one star and delta wound secondaries and rectifiers resulting in a total of

twelve pulses per cycle. These generators have very low ripple factor.

-------------------------------------------------------------------------------------------------------

-

Falling Load Generator

-------------------------------------------------------------------------------------------------------

-See the Mobile Generator notes in the Tutorials section for details of generators

used in Mobile equipment

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-Medium-frequency generator,

A state of the art generator design, also named high-frequency generator and inverter

generator, which uses a high-frequency current to produce nearly constant potential

voltage to the X ray tube with a transformer of much smaller size than found in

ordinary X-ray generators.

The incoming power supply to a medium frequency generator may be an ordinary 50

Hz (230 V) single phase current (Fig.1). This current is rectified and smoothed and

then fed to a chopper and inverter circuit which transforms the smooth, direct current

(DC) into a high-frequency (5 - 100 kHz) alternating current (AC). (The chopper

"chops" the continuous DC into high-frequency DC pulses and the inverter transforms

this into AC.) A transformer converts this high-frequency low-voltage AC into high-

voltage AC, which then is rectified by half wave rectification and smoothed to

provide a nearly constant potential high voltage to the X-ray tube. The voltage is

controlled by varying the frequency of the chopper/inverter circuit, which determines

the frequency of the current delivered to the transformer. Fast exposure switching, in

the order of 1 ms, is easily obtained with the medium frequency generator.

Outline of a medium frequency generator

(Siemens)

Waveforms from the various circuit parts

(http://www.amershamhealth.com/medcyclopaedia)

Advantages and Disadvantages

One of the great benefits of this generator design, is the reduced weight and size. The

main components of the generator may be placed within the same enclosure as the X-

ray tube, or in e.g. the C-arm of the equipment. This generator principle was

previously used only in small mobile and/or battery-powered generators with low

power rating, but today it is applied to all modern high-voltage generators up to the

highest needed power ratings above 100 kW.

Typical Generator Diagrams (Siemens)

Typical Fluoroscopy Unit Block Diagram (Siemens)

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-

Advantages and Disadvantages of different generator types

The design of a generator needs to optimise the following points, and be matched as

closely as possible to the clinical requirements of the generator usage, a point to note

is that the late 1990 saw the pinnacle of development of film screen radiography with

exposure values required for typical examinations being much less than even ten years

earlier, thus a typical chest x-ray in 1979 before rare earth screens may have been

around 25 mAS @70Kv whilst today this may have dropped to as little as 2mAS@90

Kv, interestingly some of the new Digital radiography systems require more exposure

than the film screen combinations they replace.

Efficiency of conversion of electrical energy to useful x-ray energy

Maximum dose rate per mAS

Power output

Low Ripple

Cost

Size / weight

Minimum exposure time

Reliability

Advantages and Disadvantages of Constant Potential and multiphase generators

compared with a basic single phase generator. (Stockley)

Advantages

More efficient conversion of electrical power to x-ray energy

More x-rays generated per mAS

Shorter exposure times possible

Sleeplessly variable range of exposure times

Disadvantages

More expensive to purchase

X-ray tube to cope with higher loading required

Lower image contrast

Shorter tube life

Equipment larger and heavier

Possibly more prone to failure due to greater complexity

Gradient Screens

In certain imaging situations, the attenuation of the X-rays by the human body varies

extremely within the body sections to be imaged on one and the same film. The lateral

views of the lumbar spine and of the transition between the thoracic spine and the

cervical spine are prominent examples for this dramatic variation in attenuation. It is

virtually impossible to image these areas adequately without special aids.

One way to overcome the problem caused by the extreme variation in transparency is

the use of a shaped filter with varying thickness. The thicker area of the filter is

placed in the beam where the object is more transparent. Thus, the beam intensity at

the film will be more uniform.

Another approach to achieve a more uniformly exposed image is the use of a screen

set with varying sensitivity. This type of screens is called gradient screens. Two

different basic designs are available: Either the thickness of the phosphor layers varies

across the screens, or the screens have a uniform phosphor layer thickness, but are

covered with a laquer layer of varying transparency to the light emitted by the screen.

In order to obtain a more uniformly exposed image, the less sensitive screen areas

have to be placed under the more transparent sections of the object.

Grid-controlled x-ray tube

Grid-controlled x-ray tube,

an X-ray tube which is equipped with a grid, i.e. an extra electrode between the

cathode and anode to control the flow of electrons. The third electrode is actually the

focusing cup that surrounds the filament. Normally, the focusing cup is kept at the

same negative potential as the filament. In a grid-controlled tube, the focusing cup

may be negatively charged (as compared to the filament) to such an exist that the flow

of electrons from the filament is completely stopped. The voltage applied between the

focusing cup and filament may thus act like a switch to turn the tube current on and

off. This is particularly useful when very short exposures are needed, e.g. in

cinefluorography. Grid-controlled X-ray tubes provide secondary switching as

opposed to the primary switching provided by e.g.the silicon controlled rectifier SCR .

See also exposure switching

Gurney – Mott hypothesis

The theory concerning the formation of electron traps in a silver halide crystal which

is exposed to light or X-rays. The traps can capture electrons released by ionization in

the crystal. These negatively charged electron traps which are produced during film

exposure, can attract the interstitial silver ions in the crystal. The silver ions are then

reduced to atomic silver in the traps. These few silver atoms present in a silver halide

grain following exposure act as catalyst in the development process, so that the rest of

the silver ions in the grain are reduced to metallic silver.

Half-Value Layer

A simple and commonly accepted way of characterizing the hardness or penetrating

power of an X-ray beam is determining and stating the thickness of aluminum

filtration required to cut the intensity of the beam in half.

This filtration is called the half-value layer (HVL). The higher the half-value layer is,

the harder is the beam. The penetration power increases with increasing half value

layer, but at the same time the achievable contrast decreases. The latter, however, is

not true, if edge filtration is used, as then increased filtration by the edge filter will

increase the half-value layer of the primary beam hitting the object, but at the same

time narrow the spectrum and thus improve the contrast in the image.

The half-value layer can also be used for an indirect determination of the total

filtration an X-ray beam has been subjected to (cf. aluminum equivalent).

The actual value of the filter thickness will depend somewhat on the type of

aluminum used. Pure aluminum will yield slightly higher values than aluminum alloys,

e.g. 1100 aluminum, which usually contain some copper. For the spectra used in

general radiography, this difference is practically irrelevant, but in mammography

with acceleration voltages between approximately 25 and 30 kV, the results differ

appreciably and the aluminum type used is of importance.

Introduction to x-ray film

INTRODUCTION

Photographic film can be exposed directly to X-rays but its sensitivity is very low and

prohibitively large patient exposures would result if this appraoch was implemented

on its own. Therefore, almost all conventional radiographic examinations require that

the image be converted to light by an intensifying screen before being recorded by the

film.

We will consider pertinent features of both Intensifying Screens and X-ray film below.

FLUORESCENCE

We have seen previously that luminescence refers to the stimulated (by light, ionising

radiation, chemical reactions etc.) emission of light by certain materials. If the light is

emitted instantaneously, that is within 10 nanoseconds, the phenomenon is called

fluorescence. If the emission is delayed somewhat, it is called

phosphorescence. More particularly, in radiology, fluorescence is the term used to

describe the ability of certain inorganic phosphors to emit light when excited by X-

rays.

Until the early 1970s the only phosphor of note was calcium tungstate (CaWO4), but

since then a plethora of rare-earth phosphors with improved efficiency have appeared

on the scene. No matter what type of phosphor material is used, the conversion of a

relatively small number of X-ray photons of high energy to a large number of light

photons of low energy is due predominantly to X-ray absorption via the photoelectric

effect in the high Z components of the phosphor.

The incident X-ray photons are absorbed either totally or partially in the phosphor

layer. The absorbed energy is transferred to electrons which in turn deposit their

energy by ionisation and excitation. The energy added to the atoms of the phosphor

raises the atomic electrons to excited states. Most of this added energy is then

dissipated as heat but a fraction (5% - 20%) is radiated as electromagnetic radiation in

the visible or near visible wavelengths and it is this radiation which is utilised in the

production of the latent image on the X-ray film.

INTENSIFYING SCREENS

The use of intensifying screens has three major benefits:

o Reduction of patient dose

o Reduction of tube and generator loading and

o Reduction of patient motion artifacts.

However, there is one disadvantage that is occasionally relevant to radiology which is

that the image clarity is degraded in comparison with a directly exposed film.

Figure 1 gives a schematic of a typical screen. The thin protective layer provides

protection for the phosphor and can easily be cleaned. In some screens, the reflecting

layer is not included. In a typical situation, two screens are used, one on either side of

a double emulsion film To compensate for the absorption of some X-rays by the front

screen, the back screen may be thicker than the front screen.

Figure 1: Cross-section of a typical intensifying screen. 1 micron = 1 mm.

The isotropic emission and scattering of light photons in the phosphor results in the

lateral diffusion of the scintillation pulse before it escapes the screen. This results in a

loss of resolution or sharpness and becomes increasingly important as the screen

thickness is increased. This can be compensated for by using light absorbing dyes in

the screen which will preferentially absorb the photons that travel the greatest

distances.

RARE EARTH SCREENS

We have already noted that the interaction of diagnostic X-rays with screens occurs

primarily via the photoelectric effect. Therefore we can say that we need our

phosphors to have K-edges appropriately matched to the X-ray photon

energies. More explicitly, this means that we want a phosphor whose K-edge is

between 25 and 50 keV.

You may recall that the photoelectric effect interaction probability is a maximum at

energies just above the K-edge. A look at Figure 2 establishes that Gd2O2S has a

significant advantage over calcium tungstate for photon energies between 50 and 70

keV. The same is true of other rare-earth type screens such as BaSrSO4 to a slightly

lesser extent. It is also useful to note that Gd-based phosphor screens are more

favourably disposed to the detection of primary radiation than scatter radiation as a

greater proportion of the primary spectrum is above the K-edge of Gd than of the

scatter spectrum.

Figure 2: Approximate Screen Absorption as a Function of Photon Energy for pairs of CaWO4, Gd2O2S

and BaSrSO4 screens. The spectrum from an X-ray tube operated at 80 kVp with 12.5 cm of perspex as

phantom is also illustrated.

Most inorganic phosphors (calcium tungstate is an exception) do not emit light

efficiently unless doped with a small quantity of activator. For example, the activator

in the rare-earth oxysulphides is terbium (Tb). The concentration of the activator not

only affects the amount of light emitted but the spectral emission as well. This can be

used to advantage to achieve better spectral matching between the phosphor and the

film response. Certainly, the use of these activators is the reason for the substantially

improved conversion efficiency of the rare-earth screens compared with the old

calcium tungstate screens.

X-RAY FILM

The major recording medium used in radiology is X-ray film - although the situation

is changing with the introduction of new technologies in recent years. The film can

be exposed by the direct action of X-rays, but more commonly the X-ray energy is

converted into light by intensifying screens and this light is used to expose the film, as

described above. The basic structure of the film is outlined in Figure 3 below.

Figure 3: Cross-section through a double emulsion film

The film base provides the structural strength for the film. However, the base must

be flexible for ease of processing, essentially be transparent to light and be

dimensionally stable over time. Early base materials were glass and cellulose nitrate,

but more recently cellulose triacetate and polyester have been adopted. A thin layer

of adhesive is then applied to the base and this binds the emulsion layer. Covering

the emulsion is a thin supercoat that serves to protect the emulsion from mechanical

damage.

The two most important ingredients of a photographic emulsion are gelatin and silver

halide. With most X-ray film the emulsion is coated on both sides of the film but its

thickness varies with the nature and type of the film, but is usually no thicker than 10

mm. Photographic gelatin is made from bone and is ideal as a suspension medium in

that it prevents clumping of grains. In addition, processing chemicals can penetrate

gelatin rapidly without destroying its strength or permanence.

Silver halide is the light sensitive material in the emulsion. In X-ray film, sensitivity

is increased by having a mixture of between 1% and 10% silver iodide and 90 to 99%

silver bromide. In photographic emulsion the silver halide is suspended in the gelatin

as small crystals (called grains). Grain size might average one to 2.3 mm in diameter

with up to a billion silver ions per grain and billions of grains per ml of emulsion. In

its pure form the silver halide crystal has low photographic sensitivity. The emulsion

is sensitised by heating it under controlled conditions with a reducing agent

containing sulphur. This results in the production of silver sulphide at a site on the

surface of the crystal referred to as a sensitivity speck. It is the sensitivity speck that

traps electrons to begin formation of the latent image centres.

Silver bromide is cream coloured and absorbs ultraviolet and blue light, but reflects

green and red light. Historically, this was fine since the principle emission from

calcium tungstate screens is blue light. Films for photography of image intensifier

images and films for use with rare earth screens need to have their spectral sensitivity

broadened to encompass the longer wavelengths associated with the emissions from

these screens. This is accomplished by the addition of suitable dyes. Thus, we have

green sensitive orthochromatic film and red sensitive panchromatic film.

FILM PROCESSING

Film processing is a multi-stage process involving development, fixing, washing and

replenishment (Figure 4). In development, the exposed grains are preferentially

reduced to black metallic silver. In fixing the remaining unexposed grains are

dissolved so that they can be removed from the emulsion by washing. Replenishment

ensures that chemical balance is maintained with usage of the processing solutions.

Figure 4: Schematic of an automatic film processor, showing the pathway followed by film as it is

guided by roller mechanisms through the processing solutions.

PHOTOGRAPHIC CHARACTERISTICS OF X-RAY FILM

When the X-ray beam passes through body tissues, variable fractions of the beam will

be absorbed, depending on the composition and thickness of the tissues and the

quality (kVp & filtration) of the beam. The magnitude of this variation in intensity is

the mechanism by which the X-ray beam emanating from the patient produces

diagnostic information. The information content of this X-ray image must be

transformed into a visible image on the X-ray film with minimal information loss.

In general radiography, the X-ray image is first converted to a light image using

intensifying screens, which in turn produce a visible pattern of metallic black silver on

the X-ray film. Ultimately, the degree of blackening is related to the intensity of the

radiation reaching the intensifying screen. The amount of blackness on the film is

called the optical density, D, which is defined in Figure 5. For example, if 100 light

photons are incident on a film and only one is transmitted the film density would be

log10(100) or 2. Useful densities in diagnostic radiology range from about 0.2 to

about 2.5. High density means black films.

Figure 5: The definition of optical density, D.

If the relationship between the logarithm of the radiation exposure and the optical

density is plotted we obtain a curve known as the Characteristic Curve. For film

exposed with an intensifying screen, this curve is essentially sigmoidal in shape

(Figure 6). It is characterised by:

o a toe or region of low gradient at low exposures,

o a region of relatively steep increase in density for minimal exposure

increases, and

o a third relatively flat region called the shoulder at high exposures.

The important part of the curve diagnostically is the approximately linear region

between the toe and the shoulder where the density is proportional to the logarithm of

the exposure.

Figure 6: The Characteristic Curve of X-ray film.

The information content resulting from the radiograph arises from differences in the

film density, which we can define as radiographic contrast. Radiographic contrast

depends on subject contrast and film contrast. For the moment you should recall

that subject contrast depends on the differential attenuation of the X-ray flux as it

passes through the patient and is affected by thickness, density and atomic number of

the irradiated parts of the subject, the kVp, the presence of contrast medium and

scattered radiation. For example, relatively few X-ray photons pass through bone

compared with soft tissue but care must be taken in selecting the correct kVp in order

to produce an X-ray image of high information content for the screen-film to

record. That is, the kVp influences the magnitude of the subject contrast.

Film contrast depends on four factors:

o the characteristic curve of the film,

o the film density,

o use of intensifying screens or direct exposure and

o the film processing.

The slope of the straight line portion of the characteristic curve tells us how much

change in film density will occur as exposure changes. The slope or gradient of the

curve may be measured and the maximum gradient is called the film gamma, which

tells us how well the film will amplify the subject contrast.

X-ray film will fog slowly with time, the extent depending markedly on how well it is

stored. This fogging, along with the optical density of the film base, will generate a

low density in the toe section of the Characteristic Curve.

The shoulder region of the curve indicates over exposure

Ion

Definition

Ion, any atom or electron which has a positive or negative electric charge owing to an

electrical imbalance between its atomic protons and electrons

Ionization

The removal of an electron from an atom causing the electrical balance to be net

positive

Newton's Inverse Square Law

Any point source which spreads its influence equally in all directions without a limit

to its range will obey the inverse square law. This comes from strictly geometrical

considerations. The intensity of the influence at any given radius (r) is the source

strength divided by the area of the sphere. Being strictly geometric in its origin, the

inverse square law applies to diverse phenomena. Point sources of gravitational force,

electric field, light, sound, or radiation obey the inverse square law.

As one of the fields which obey the general inverse square law, a point radiation

source can be characterized by the diagram above whether you are talking about

Roentgens, rads, or rems. All measures of exposure will drop off by the inverse

square law. For example, if the radiation exposure is 100 mR/hr at 1 inch from a

source, the exposure will be 0.01 mR/hr at 100 inches.

To calculate a new exposure maS ie one at a new distance using the old exposure maS

New mAs = Old mAs x (New distance2/Old distance

2)

eg A Chest x-ray at 180 cm and using 5 mAS

What mAS would be needed supine on a trolley at 100cm?

New mAS = 5 [old mAS] x (10000 [New d2] / 32400 [Old d

2]

New mAS = 5 x [10000 / 32400] 0.3 = 1.5 mAS

Optical Density

The relative darkness of an image in a finished film is called simply density.

Photographic density for film is determined by measuring the incident to transmitted

light ratio and expressing the value as a logarithm

Density (D) = Log (Intensity of Incident Light / Intensity of Transmitted Light)

For example if 100 is the intesity of the viewing bow and 10 is the value transmitted

through a film the film density = Log (100/10) =1

Photostimulable phosphor plate

A radiographic screen containing a special class of phosphors which when exposed to

X-rays, stores the latent image as a distribution of electron charges, the energy of

which may later be freed as light by stimulation with a scanning laser beam. The light

is directed to a photomultiplier tube, and the output electrical signal is digitized. The

final result is a digital projection radiograph. The photostimulable phosphor plate is

also known as an imaging plate, storage phosphor imaging plate, and digital cassette.

The technique has also been termed computed radiography (CR) (after the

introduction of the imaging plate in 1981 by the Fuji company, who named the new

technique FCR).

The photostimulable phosphors in the imaging plate have a property termed

phosphorescence or photoluminescence (see luminescent screen) which in this context

means they are able to store X-ray energy and later, when stimulated by (laser)light,

free the energy as emitted light. The phosphors used in radiography are mixtures of

three different barium fluorohalides doped with europium as an activator; BaFI:Eu2+,

BaFCl:Eu2+, and BaFBr:Eu2+. To prepare the imaging plate for an X-ray exposure,

the plate is exposed to intense light to erase any previous image. For X-ray imaging,

the plate is placed in a cassette and is used just like a film screen cassette with

standard radiographic euipment. When exposed to X-rays, the europium atoms in the

phosphor crystalline lattice are ionized (converted from 2+ to 3+), liberating a valence

electron. These electrons are raised to a higher energy state in the conduction band

(see solid and photoconduction for an explanation of conduction band). Once in the

conduction band, the electrons travel freely until they are trapped in a so-called F-

centre in a metastable state with an energy level slightly below that of the conduction

band, but higher than that of the valence band. The number of trapped electrons is

proportional to the amount of X-rays absorbed locally. The trapped electrons

constitute the latent image. Due to thermal motion, the electrons will slowly be

liberated from the traps, and the latent image should therefore be read without too

much delay. At room temperature, the image should, however, be readable up to 8

hours after exposure.

Reading of the exposed imaging plate is performed by scanning the plate with a small

(50–200 mm) dot of light from a helium-neon laser. The laser light stimulates the

trapped electrons up to the conduction band, where they are free to move to the

europium atoms, thereby leaving the high energy conduction band to return to the

lower energy valence band. The transformation of europium from the 3+ to the 2+

state therefore involves liberation of energy, and this is done by emission of light.

Since there is a larger energy difference between the conduction band and the valence

band than between the conduction band and the F-centres, the (green) light emitted

has a higher energy than the (red) laser light needed to stimulate the trapped electrons.

The difference in wavelength between the two lights is critical for detection of the

emitted light. By using a filter that absorbs red light but is transparent to green light,

the emitted light is selectively detected. The laser beam scans the imaging plate in a

transverse direction while the plate is moved past the scanning beam. The emitted

light is collected using a light guide and is fed to a photomultiplier tube where the

light is converted to an electrical signal which is amplified to an electric output signal.

This signal is digitized, and the image is stored in a computer as a digital matrix, each

pixel having a gray scale value determined by the amount of light emitted from the

corresponding dot on the imaging plate.

The imaging plate has a much wider dynamic range than film-screen systems, with a

linear characteristic curve (see digital radiography (I), Fig. 2), giving the system a

much wider exposure latitude than film-screen systems. Because of certain pre-scan

operations performed prior to the actual read-out of the imaging plate, an automatic

gain control is achieved; overexposed images are recorded with equal "brightness" as

underexposed images. The required amount of radiation to the plate is, however, in

average the same as needed with film-screen systems. Due to the wide exposure

latitude and "automatic gain control", doses may be reduced, but at the cost of

increased noise. The uniform density despite over- and underexposure is one of the

great benefits of the system as compared to conventional film-screen systems; almost

no retakes due to incorrect exposures are necessary. Additional benefits are those

common to all digital techniques, including postprocessing such as changing window

level and width, exact measurement of distances, angles, and areas, zooming, panning,

and not the least, digital archiving and communication (see PACS).

Production and Properties of X Rays

X rays are produced when fast moving electrons hit a piece of metal (called the

target).

Electrons are thermionically emitted by the filament (cathode).

The accelerating voltage is about 100kV.

Less than 1% of the kinetic energy of the electrons is converted into x rays so the

anode (target) must be cooled during operation.

X rays are not deflected by electric or magnetic fields but can be diffracted suggesting

that they have wave-like properties.

X rays are electro-magnetic radiations having wavelengths in the range 10-11

m to 10-

8m.

X rays cause certain substances to fluoresce, they affect photo-graphic emulsions

and can ionise atoms. These three properties can be used to detect x rays.

The intensity of the beam of x rays (Wm-2

) depends on the number of electrons

hitting the target per unit time. This depends on the temperature of the filament.

The penetrating power of the beam of x rays depends on the kinetic energy of the

electrons. This depends on the accelerating voltage.

Quality & intensity of an X-Ray Beam

The intensity of the beam of x rays (Wm-2

) depends on the number of electrons

hitting the target per unit time ie the number of electrons flowing through the

tube the tube current. This depends on the temperature of the filament.

The penetrating power of the beam of x rays depends on the kinetic energy of the

electrons. This depends on the accelerating voltage.

A useful analogy of this can be made with light. Intensity can be equated to colour and penetrating power to brightness.

Increase of the accelerating voltage applied between filament and target is found to

increase the penetrating power of the Xrays. Since the maximum loss of kinetic

energy at a single collision is now higher (=eV), the highest frequency emitted is also

higher as expected. Thus the quality of the emitted X rays is altered. These are called

‗hard‘ Xrays.

It is found that an increase of the heating voltage increases the intensity of Xrays

without any change in the hardness or penetrating power. The high intensity spikes

characteristic of the target material are also of unchanged wavelengths

Relative film speed test

Split phantom test and how is it used to test sensitometric differences between

different emulsions or batches of film? (From the Kodak website)

A split phantom test should be performed to radiographically determine relative speed

differences between two different boxes of film, one of which is suspected of being

much faster or slower than the film in current use for either clinical films or for

processor quality control. Speed comparisons made using a sensitometer may not

accurately reflect the differences in speed between two films exposed by light from an

intensifying screen.

The procedure is as follows:

1.

2. Assemble the tools that are needed for the test:

o A phantom used for mammography quality control testing

o The 18 x 24 cm mammography cassette normally used for the phantom

test

o A piece of cardboard from the film box cut in half to use as a guide

o A pair of scissors

o A lead pencil

The mammography x-ray unit and the processor will also be used for this test.

3. In the darkroom (in total darkness to reduce any additional density added to

the films due to long safelight exposure) cut a sheet of film from the current or

"normal" box in half by using the cardboard as a guide. (This can be done by

lining up the 18 cm edges of the cardboard and film so that the film is closest

to the countertop and the cardboard half is on top. Be careful cutting the film

in the dark.)

4. Place the film--emulsion side up--in the cover of the opened cassette with the

film on the right side and the cut edge toward the right edge of the cassette;

use a lead pencil to mark the corner "N" for normal.

5. Cut a sheet of film from the "suspect" box in half by using the cardboard as a

guide.

6. Place the film--emulsion side up--in the cover of the opened cassette with the

film on the left side and the cut edge toward the left edge of the cassette; use

the lead pencil to mark the corner "S" for suspect.

7. Before closing the cassette, make sure the film edges in the center of the

cassette are directly adjacent to one another and not overlapping.

8. Place the cassette with the two film halves in the grid of the mammography x-

ray unit.

9. Place the phantom on top of the grid in the standard location used for

mammography quality control testing.

10. Position the photocell beneath the center of the phantom (standard location),

assuming the phantom exposure is always made using the phototimer.

11. Select the same technique factors usually employed when imaging the

phantom (same kVp, etc.).

12. Make the exposure and immediately process the two film halves in the same

manner (e.g., emulsion side up and on the right side of the processor).

13. Use a densitometer to take two optical density readings in the center of the

phantom, just to the right and left of the cut edges (one on the "normal" and

one on the "suspect" film).

14. Calculate the density difference by subtracting the optical density value of the

"suspect" film from the optical density value of the "normal" film.

If the density difference is a negative value and the "suspect" film is darker

than the "normal" film, the "suspect" film is faster. If the density difference is

a positive value and the "suspect" film is lighter than the "normal" film, the

"suspect" film is slower.

According to the American College of Radiology in Recommended Specifications for

New Mammography Equipment :

"A density difference of 0.30 between any two films of the same type from the

same manufacturer, exposed and processed together, is a reasonable maximum

to be expected from manufacturing variability for films of roughly the same

age and storage conditions."

"If the difference between the two film densities exceeds 0.30 at a density of

approximately 1.25, then the film supplier should be contacted to determine

the source of the problem."

Note that a difference of 0.30 at a density of approximately 1.25 may translate

into a bigger difference for clinical films exposed at a greater optical density. For

example, high-contrast mammography films, such as KODAK MIN-R 2000 Film,

are frequently exposed at an optical density between 1.50 and 2.00 in order to

maximize contrast. The density difference at this optical density level will be

greater due to the increased contrast.

The Photographic Latent Image

As shown in earlier figures, a photographic emulsion consists of a myriad of tiny

crystals of silver halide--usually the bromide with a small quantity of iodide--

dispersed in gelatin and coated on a support. The crystals--or photographic grains--

respond as individual units to the successive actions of radiation and the photographic

developer.

The photographic latent image may be defined as that radiation-induced change in a

grain or crystal that renders the grain readily susceptible to the chemical action of a

developer.

To discuss the latent image in the confines of this siterequires that only the basic

concept be outlined. A discussion of the historical development of the subject and a

consideration of most of the experimental evidence supporting these theories must be

omitted because of lack of space.

It is interesting to note that throughout the greater part of the history of photography,

the nature of the latent image was unknown or in considerable doubt. The first public

announcement of Daguerre's process was made in 1839, but it was not until 1938 that

a reasonably satisfactory and coherent theory of the formation of the photographic

latent image was proposed. That theory has been undergoing refinement and

modification ever since.

Some of the investigational difficulties arose because the formation of the latent

image is a very subtle change in the silver halide grain. It involves the absorption of

only one or a few photons of radiation and can therefore affect only a few atoms, out

of some 109 or 1010 atoms in a typical photographic grain. The latent image cannot be

detected by direct physical or analytical chemical means.

However, even during the time that the mechanism of formation of the latent image

was a subject for speculation, a good deal was known about its physical nature. It was

known, for example, that the latent image was localized at certain discrete sites on the

silver halide grain. If a photographic emulsion is exposed to light, developed briefly,

fixed, and then examined under a microscope (see the figure below), it can be seen

that development (the reduction of silver halide to metallic silver) has begun at only

one or a few places on the crystal. Since small amounts of silver sulfide on the surface

of the grain were known to be necessary for a photographic material to have a high

sensitivity, it seemed likely that the spots at which the latent image was localized

were local concentrations of silver sulfide.

Electron micrograph of exposed, partially developed, and fixed grains, showing

initiation of development at localized sites on the grains

(1µ = 1 micron = 0.001 mm).

It was further known that the material of the latent image was, in all probability, silver.

For one thing, chemical reactions that will oxidize silver will also destroy the latent

image. For another, it is a common observation that photographic materials given

prolonged exposure to light darken spontaneously, without the need for development.

This darkening is known as the print-out image. The printout image contains enough

material to be identified chemically, and this material is metallic silver. By

microscopic examination, the silver of the print-out image is discovered to be

localized at certain discrete areas of the grain (see the figure below), just as is the

latent image.

Electron micrograph of photolytic silver produced in a grain by very intense

exposure to light.

Thus, the change that makes an exposed photographic grain capable of being

transformed into metallic silver by the mild reducing action of a photographic

developer is a concentration of silver atoms--probably only a few--at one or more

discrete sites on the grain. Any theory of latent-image formation must account for the

way that light photons absorbed at random within the grain can produce these isolated

aggregates of silver atoms. Most current theories of latent-image formation are

modifications of the mechanism proposed by R. W. Gurney and N. F. Mott in 1938.

In order to understand the Gurney-Mott theory of the latent image, it is necessary to

digress and consider the structure of crystals--in particular, the structure of silver

bromide crystals.

When solid silver bromide is formed, as in the preparation of a photographic emulsion,

the silver atoms each give up one orbital electron to a bromine atom. The silver atoms,

lacking one negative charge, have an effective positive charge and are known as silver

ions (Ag+). The bromine atoms, on the other hand, have gained an electron--a negative

charge--and have become bromine ions (Br-). The "plus" and "minus" signs indicate,

respectively, one fewer or one more electron than the number required for electrical

neutrality of the atom.

A crystal of silver bromide is a regular cubical array of silver and bromide ions, as

shown schematically in the figure below. It should be emphasized that the

"magnification" of the figure is very great. An average grain in an industrial x-ray

film may be about 0.00004 inch in diameter, yet will contain several billions of ions.

A silver bromide crystal is a rectangular array of silver (Ag+) and bromide (Br-)

ions.

A crystal of silver bromide in a photographic emulsion is--fortunately--not perfect; a

number of imperfections are always present. First, within the crystal, there are silver

ions that do not occupy the "lattice position" shown in the figure above, but rather are

in the spaces between. These are known as interstitial silver ions (see the figure

below). The number of the interstitial silver ions is, of course, small compared to the

total number of silver ions in the crystal. In addition, there are distortions of the

uniform crystal structure. These may be "foreign" molecules, within or on the crystal,

produced by reactions with the components of the gelatin, or distortions or

dislocations of the regular array of ions shown in the figure above. These may be

classed together and called "latent-images sites."

"Plain view" of a layer of ions of a crystal similar to that of the previous figure.

A latent-image site is shown schematically, and two interstitial silver ions are

indicated.

The Gurney-Mott theory envisions latent-image formation as a two-stage process. It

will be discussed first in terms of the formation of the latent image by light, and then

the special considerations of direct x-ray or lead foil screen exposures will be covered

THE GURNEY-MOTT THEORY

When a photon of light of energy greater than a certain minimum value (that is, of

wavelength less than a certain maximum) is absorbed in a silver bromide crystal, it

releases an electron from a bromide (Br-) ion. The ion, having lost its excess negative

charge, is changed to a bromine atom. The liberated electron is free to wander about

the crystal (see the figure below). As it does, it may encounter a latent image site and

be "trapped" there, giving the latent-image site a negative electrical charge. This first

stage of latent-image formation--involving as it does transfer of electrical charges by

means of moving electrons--is the electronic conduction stage.

Stages in the development of the latent image according to the Gurney-Mott

theory.

The negatively charged trap can then attract an interstitial silver ion because the silver

ion is charged positively (C in the figure above). When such an interstitial ion reaches

a negatively charged trap, its charge is counteracted, an atom of silver is deposited at

the trap, and the trap is "reset" (D in the figure above). This second stage of the

Gurney-Mott mechanism is termed the ionic condition stage, since electrical charge is

transferred through the crystal by the movement of ions--that is, charged atoms. The

whole cycle can recur several, or many, times at a single trap, each cycle involving

absorption of one photon and addition of one silver atom to the aggregate. (See E to H

in the figure above.)

In other words, this aggregate of silver atoms is the latent image. The presence of

these few atoms at a single latent-image site makes the whole grain susceptible to the

reducing action of the developer. In the most sensitive emulsions, the number of silver

atoms required may be less than ten.

The mark of the success of a theory is its ability to provide an understanding of

previously inexplicable phenomena. The Gurney-Mott theory and those derived from

it have been notably successful in explaining a number of photographic effects. One

of these effects--reciprocity-law failure--will be considered here as an illustration.

Low-intensity reciprocity-law failure (left branch of the curve ) results from the fact

that several atoms of silver are required to produce a stable latent image. A single

atom of silver at a latent-image site (D in the figure above) is relatively unstable,

breaking down rather easily into an electron and a positive silver ion. Thus, if there is

a long interval between the formation of the first silver atom and the arrival of the

second conduction electron (E in the figure above), the first silver atom may have

broken down, with the net result that the energy of the light photon that produced it

has been wasted. Therefore, increasing light intensity from very low to higher values

increases the efficiency, as shown by the downward trend of the left-hand branch of

the curve, as intensity increases.

High-intensity reciprocity-law failure (right branch of the curve) is frequently a

consequence of the sluggishness of the ionic process in latent-image formation (see

the figure above). According to the Gurney-Mott mechanism, a trapped electron must

be neutralized by the movement of an interstitial silver ion to that spot (D in the figure

above) before a second electron can be trapped there (E in the figure above);

otherwise, the second electron is repelled and may be trapped elsewhere. Therefore, if

electrons arrive at a particular sensitivity center faster than the ions can migrate to the

center, some electrons are repelled, and the center does not build up with maximum

efficiency. Electrons thus denied access to the same traps may be trapped at others,

and the latent image silver therefore tends to be inefficiently divided among several

latent-image sites. (This has been demonstrated by experiments that have shown that

high-intensity exposure produces more latent image within the volume of the crystal

than do either low- or optimum-intensity exposures.) Thus, the resulting inefficiency

in the use of the conduction electrons is responsible for the upward trend of the right-

hand branch of the curve.

X-RAY LATENT IMAGE

In industrial radiography, the photographic effects of x-rays and gamma rays, rather

than those of light, are of the greater interest.

At the outset it should be stated that the agent that actually exposes a photographic

grain, that is, a silver bromide crystal in the emulsion, is not the x-ray photon itself,

but rather the electrons--photoelectric and Compton--resulting from the absorption

event. It is for this reason that direct x-ray exposures and lead foil screen exposures

are similar and can be considered together.

The most striking differences between x-ray and visible-light exposures to grains arise

from the difference in the amounts of energy involved. The absorption of a single

photon of light transfers a very small amount of energy to the crystal. This is only

enough energy to free a single electron from a bromide (Br-) ion, and several

successive light photons are required to render a single grain developable. The

passage through a grain of an electron, arising from the absorption of an x-ray photon,

can transmit hundreds of times more energy to the grain than does the absorption of a

light photon. Even though this energy is used rather inefficiently, in general the

amount is sufficient to render the grain traversed developable--that is, to produce

within it, or on it, a stable latent image.

As a matter of fact, the photoelectric or Compton electron, resulting from absorption

or interaction of a photon, can have a fairly long path in the emulsion and can render

several or many grains developable. The number of grains exposed per photon

interaction can vary from 1 grain for x-radiation of about 10 keV to possibly 50 or

more grains for a 1 meV photon. However, for 1 meV and higher energy photons,

there is a low probability of an interaction that transfers the total energy to grains in

an emulsion. Most commonly, high photon energy is imparted to several electrons by

successive Compton interactions. Also, high-energy electrons pass out of an emulsion

before all of their energy is dissipated. For these reasons there are, on the average, 5

to 10 grains made developable per photon interaction at high energy.

For comparatively low values of exposure, each increment of exposure renders on the

average the same number of grains developable, which, in turn, means that a curve of

net density versus exposure is a straight line passing through the origin (see the figure

below). This curve departs significantly from linearity only when the exposure

becomes so great that appreciable energy is wasted on grains that have already been

exposed. For commercially available fine-grain x-ray films, for example, the density

versus exposure curve may be essentially linear up to densities of 2.0 or even higher.

Typical net density versus exposure curves for direct x-ray exposures.

The fairly extensive straight-line relation between exposure and density is of

considerable use in photographic monitoring of radiation, permitting a saving of time

in the interpretation of densities observed on personnel monitoring films.

It the D versus E curves shown in the figure above are replotted as characteristic

curves (D versus log E), both characteristic curves are the same shape (see the figure

below) and are merely separated along the log exposure axis. This similarity in toe

shape has been experimentally observed for conventional processing of many

commercial photographic materials, both x-ray films and others.

Characteristic curves plotted from the data in the previous figure.

Because a grain is completely exposed by the passage of an energetic electron, all x-

ray exposures are, as far as the individual grain is concerned, extremely short. The

actual time that an x-ray-induced electron is within a grain depends on the electron

velocity, the grain dimensions, and the "squareness" of the hit. However, a time of the

order of 10-13 second is representative. (This is in distinction to the case of light where

the "exposure time" for a single grain is the interval between the arrival of the first

photon and that of the last photon required to produce a stable latent image.)

The complete exposure of a grain by a single event and in a very short time implies

that there should be no reciprocity-law failure for direct x-ray exposures or for

exposures made with lead foil screens. The validity of this has been established for

commercially available film and conventional processing over an extremely wide

range of x-ray intensities. That films can satisfactorily integrate x-, gamma-, and beta-

ray exposures delivered at a wide range of intensities is one of the advantages of film

as a radiation dosimeter.

In the discussion on reciprocity-law failure it was pointed out that a very short, very

high intensity exposure to light tends to produce latent images in the interior of the

grain. Because x-ray exposures are also, in effect, very short, very high intensity

exposures, they too tend to produce internal, as well as surface, latent images.

DEVELOPMENT

Many materials discolor on exposure to light--a pine board or the human skin, for

example--and thus could conceivably be used to record images. However, most such

systems reset to exposure on a "1:1" basis, in that one photon of light results in the

production of one altered molecule or atom. The process of development constitutes

one of the major advantages of the silver halide system of photography. In this system,

a few atoms of photolytically deposited silver can, by development, be made to

trigger the subsequent chemical deposition of some 109 or 1010 additional silver atoms,

resulting in an amplification factor of the order of 109 or greater. The amplification

process can be performed at a time, and to a degree, convenient to the user and, with

sufficient care, can be uniform and reproducible enough for the purposes of

quantitative measurements of radiation.

Development is essentially a chemical reduction in which silver halide is reduced or

converted to metallic silver in order to retain the photographic image, however, the

reaction must be limited largely to those grains that contain a latent image. That is, to

those grains that have received more than a certain minimum exposure to radiation.

Compounds that can be used as photographic developing agents, therefore, are limited

to those in which the reduction of silver halide to metallic silver is catalyzed (or

speeded up) by the presence of the metallic silver of the latent image. Those

compounds that reduce silver halide in the absence of a catalytic effect by the latent

image are not suitable developing agents because they produce a uniform overall

density on the processed film.

Many practical developing agents are relatively simple organic compounds (see the

figure below) and, as shown, their activity is strongly dependent on molecular

structure as well as on composition. There exist empirical rules by which the

developing activity of a particular compound may often be predicted from a

knowledge of its structure.

Configurations of dihydroxybenzene, showing how developer properties depend

on structure.

The simplest concept of the role of the latent image in development is that it acts

merely as an electron-conducting bridge by which electrons from the developing

agent can reach the silver ions on the interior face of the latent image. Experiment has

shown that this simple concept is inadequate to explain the phenomena encountered in

practical photographic development. Adsorption of the developing agent to the silver

halide or at the silver-silver halide interface has been shown to be very important in

determining the rate of direct, or chemical, development by most developing agents.

The rate of development by hydroquinone (see the figure above), for example,

appears to be relatively independent of the area of the silver surface and instead to be

governed by the extent of the silver-silver halide interface.

The exact mechanisms by which a developing agent acts are relatively complicated,

and research on the subject is very active.

The broad outlines, however, are relatively clear. A molecule of a developing agent

can easily give an electron to an exposed silver bromide grain (that is, to one that

carries a latent image), but not to an unexposed grain. This electron can combine with

a silver (Ag+) ion of the crystal, neutralizing the positive charge and producing an

atom of silver. The process can be repeated many times until all the billions of silver

ions in a photographic grain have been turned into metallic silver.

The development process has both similarities to, and differences from, the process of

latent-image formation. Both involve the union of a silver ion and an electron to

produce an atom of metallic silver. In latent image formation, the electron is freed by

the action of radiation and combines with an interstitial silver ion. In the development

process, the electrons are supplied by a chemical electron-donor and combine with the

silver ions of the crystal lattice.

The physical shape of the developed silver need have little relation to the shape of the

silver halide grain from which it was derived. Very often the metallic silver has a

tangled, filamentary form, the outer boundaries of which can extend far beyond the

limits of the original silver halide grain (see the figure below). The mechanism by

which these filaments are formed is still in doubt although it is probably associated

with that by which filamentary silver can be produced by vacuum deposition of the

silver atoms from the vapor phase onto suitable nuclei.

Electron micrograph of a developed silver bromide grain.

The discussion of development has thus far been limited to the action of the

developing agent alone. However, a practical photographic developer solution

consists of much more than a mere water solution of a developing agent. The function

of the other common components of a practical developer are the following:

An Alkali

The activity of developing agents depends on the alkalinity of the solution. The alkali

should also have a strong buffering action to counteract the liberation of hydrogen

ions--that is, a tendency toward acidity--that accompanies the development process.

Common alkalis are sodium hydroxide, sodium carbonate, and certain borates.

A Preservative

This is usually a sulfite. One of its chief functions is to protect the developing agent

from oxidation by air. It destroys certain reaction products of the oxidation of the

developing agent that tend to catalyze the oxidation reaction. Sulfite also reacts with

the reaction products of the development process itself, thus tending to maintain the

development rate and to prevent staining of the photographic layer.

A Restrainer

A bromide, usually potassium bromide, is a common restrainer or antifoggant.

Bromide ions decrease the possible concentration of silver ions in solution (by the

common-ion effect) and also, by being adsorbed to the surface of the silver bromide

grain, protect unexposed grains from the action of the developer. Both of these actions

tend to reduce the formation of fog.

Commercial developers often contain other materials in addition to those listed above.

An example would be the hardeners usually used in developers for automatic

processors.

Transmission, absorption, scatter and attenuation

Transmission

X-ray photons that pass through the patient unchanged

Absorption

X-ray photons that transfer their energy to the patient

The absorption of the X-ray radiation by a material is proportional to the degree of X-

ray attenuation and is dependent on the energy of the X-ray radiation and the

following material parameters: Thickness; Density; Atomic number

Scatter

Radiation that, during its passage through a substance, has been changed in direction.

It may also have been modified by a decrease in energy

Attenuation,

The process by which radiation loses power as it travels through matter and interacts

with it. Attenuation of x-rays in solids takes place by several different mechanisms,

some due to absorption, and others due to the scattering of the beam.

HVT

That thickness of a specified material (usually a metal) which reduces the exposure

rate to one-half its initial value.

Intensity

Relative number of x-ray photons in the x-ray beam

Quality

Quality is a measurement of the penetrating power of the X-Ray photons. The quality

of the beam increases as the proportion of high energy photons increases.

Factors affect the quality and intensity of the beam

kV - the greater the potential difference across the tube, the faster the electrons move

and the higher the energy of the X-Ray photons. Thus the quality and intensity are

increased

mA- the higher the tube current, the greater the intensities of all the photon energies,

and the intensity of the beam is increased

time - the longer the exposure, the greater the time during which X-Rays are produced,

and the greater the beam intensity. The time (seconds) and mA are generally

considered together as the composite factor mAs

distance - increasing distance from the source of radiation results in a decrease in the

intensity of the beam, according to the inverse square law. Hence doubling the

distance from the tube head will result in a beam of one quarter its original intensity.

Example values of linear coefficient of attenuation

Material Density

g cm-3

Electron

density

x 1023

g-1

Effective atomic

number

Photon

energy cm

Water 1.0 3.343 7.5 100 keV 0.17

10 MeV 0.03

Bone 1.65 3.19 12.3 100 keV 0.3

10 MeV 0.04

Lead 11.35 2.38 82 100 keV 62

10 MeV 4.3

For example, for a 100 keV x-ray beam, 1 cm of water will attenuate 17% (0.17) of

the x-ray photons in the beam.

The attenuation or absorption, usually defined as the linear absorption coefficient, µ,

is defined for a narrow well-collimated, monochromatic x-ray beam. The linear

absorption coefficient is the sum of contributions of the following:

1. Thomson scattering (R) (also known as Rayleigh, coherent, or classical

scattering) occurs when the x-ray photon interacts with the whole atom so that

the photon is scattered with no change in internal energy to the scattering atom,

nor to the x-ray photon.

2. Photoelectric (PE) absorption of x-rays occurs when the x-ray photon is

absorbed resulting in the ejection of electrons from the atom, resulting in the

ionization of the atom. Subsequently, the ionized atom returns to the neutral

state with the emission of an x-ray characteristic of the atom.

3. Compton Scattering (C) (also known as incoherent scattering) occurs when

the incident x-ray photon ejects an electron from an atom and an x-ray photon

of lower energy is scattered from the atom.

4. Pair Production (PP) can occur when the x-ray photon energy is greater than

1.02 MeV, when an electron and positron are created with the annihilation of

the x-ray photon (absorption).

5. Photodisintegration (PD) is the process by which the x-ray photon is

captured by the nucleus of the atom with the ejection of a particle from the

nucleus when all the energy of the x-ray is given to the nucleus (absorption).

This process may be neglected for the energies of x-rays used in radiography.

There are three main processes that may occur resulting in exponential

attenuation of x-ray energy:

Photoelectric absorption

Compton (inelastic scatter)

Pair production

If we compare the probability of each of these processes in water at different x-ray

photon energies, we would see something like this:

X-ray photon energy Photoelectric

absorption

Compton

scatter

Pair

production

10 keV 95% 5% 0

25 keV

(Mammography) 50% 50% 0

60 keV

(Diagnostic)

7% 93% 0

150 keV 0 100% 0

4 MeV 0 94% 6%

10 MeV

(Therapy)

0 77% 23%

24 MeV 0 50% 50%

Photoelectric absorption

Photoelectric (PE) absorption of x-rays occurs when the x-ray photon is absorbed

resulting in the ejection of electrons from the inner shell of the atom, resulting in the

ionization of the atom. Subsequently, the ionized atom returns to the neutral state with

the emission of an x-ray characteristic of the atom.

This subsequent emission of lower energy photons is generally absorbed and does not

contribute to (or hinder) the image making process. Photoelectron absorption is the

dominant process for x-ray absorption up to energies of about 500 KeV.

Photoelectron absorption is also dominant for atoms of high atomic numbers.

Photoelectric Effect is dependent on Z3 (Atomic number z)

Photoelectric absorption is a process of total absorption

Note that an ion results when the photoelectron leaves the atom.

Two subsequent points should also be noted:

Firstly, the photoelectron can cause ionisations along its track,

Secondly, X-ray emission can occur when the vacancy left by the photoelectron is

filled by an electron from an outer shell of the atom.

There are a number of rules which govern the probability of a photoelectric event:

The incident photon must have sufficient energy to overcome the binding energy of

the electron.

Once the threshold imposed by the binding energy has been exceeded, then the

interaction probability is at a maximum.

The probability of an interaction is greatest if the electron is deeply bound. That is,

the larger the atomic number, Z, the greater is the probability of a photoelectric

process

Compton Scattering

Compton Scattering, also known as incoherent scattering or inelastic scattering,

occurs when the incident x-ray photon ejects an outer shell electron from an atom and

an x-ray photon of lower energy is scattered from the atom. Relativistic energy and

momentum are conserved in this process and the scattered x-ray photon has less

energy and therefore greater wavelength than the incident x-ray photon. Compton

Scattering is important for low atomic number specimens. At energies of 100 keV --

10 MeV the absorption of radiation is mainly due to the Compton effect.

The scattered x-ray photon has an energy which is dependent on its angle of emission

and on the incident photon energy.

The probability of a Compton event depends on the number of electrons in an

absorber, which depends on the density of the absorber and the number of electrons

per unit mass. Now with the exception of hydrogen, all elements contain

approximately the same number of electrons per unit mass. Therefore the number of

Compton reactions is independent of atomic number. However, for tissues of

biological interest, the probability of an interaction does decrease slowly with

increasing photon energy above about 50 keV.

Compton scatter is an attenuation process of partial absorption and partial scatter

Pair Production (PP) is of particular importance when high-energy photons pass

through materials of a high atomic number. Energy: > 1.02 MeV

When the energy of the incident photon is greater than 1022 keV, the photon may be

absorbed through the process of Pair Production. When such a photon passes near the

nucleus of the atom it experiences the strong field of the nucleus and may be absorbed

with the creation of a positive and negative electron pair. This is an example of

conversion of energy to mass as espoused by Albert Einstein. No electronic charge is

created since the positron and electron are equal and oppositely charged. Ignoring the

tiny amount of energy given to the recoiling nucleus we may write:

E = 2mc2 + E+ + E-

where:

o m: electron rest mass,

o c: the speed of light,

o E+: kinetic energies of the positron, and

o E-: kinetic energy of the electron.

The total energy given to the electron-positron pair can be divided randomly although

there is a slightly greater probability that the positron will carry off more energy than

the electron because it experiences the repulsive Coulomb force of the nucleus'

positive charge.

The most important feature to note is that the process is not possible unless the photon

energy is greater than the rest mass energy of the electron-positron pair, i.e.

2 x 511 keV = 1022 keV.

The fate of the positron has an important bearing on the ultimate decay products. In

travelling through matter, the positron excites and ionises atoms, just as an electron

does, until it is finally brought to rest. Then it combines or annihilates with a free

electron with the production of two 511 keV photons. In order to conserve

momentum and energy the two photons move essentially at an angle of 180o to each

other.

A forth process called Coherent Scattering occurs mainly at low energies and large

values of Z and is typically a just small proportion of the total number of interaction

Here a gamma-ray or X-ray photon undergoes an interaction where it changes its

direction without loss of energy. In the idealistic situation of the interaction being

between a photon and a single free electron the process is called Thomson Scattering.

This should be contrasted with the real world situation where photons are scattered by

bound electrons. The electrons are set vibrating by the oscillating electromagnetic

field associated with the photon. Subsequently, a photon of radiation is emitted with

the same wavelength as the incident radiation leaving the atom in its original

undisturbed state. The waves from electrons within the atom combine with each other

to form the scattered wave. The scattering is a cooperative phenomenon and the

process is called Coherent Scattering. There is no net ionisation in the process, a

property which distinguishes coherent scattering from other photon interactions.

A fifth process Nuclear Photodisintegration

At extremely high energies ( > 8 MeV), a photon may interact directly with the

nucleus of an atom and eject a neutron, proton or on rare occasions even an alpha

particle.

Summary

Photoelectric (PE) absorption of x-rays occurs when the x-ray photon is

absorbed resulting in the ejection of electrons from the atom, resulting in the

ionization of the atom. Subsequently, the ionized atom returns to the neutral

state with the emission of an x-ray characteristic of the atom.

Compton Scattering (C) (also known as incoherent scattering) occurs when

the incident x-ray photon ejects an electron from an atom and an x-ray photon

of lower energy is scattered from the atom.

Pair Production (PP) can occur when the x-ray photon energy is greater than

1.02 MeV, when an electron and positron are created with the annihilation of

the x-ray photon (absorption).

Photodisintegration (PD) is the process by which the x-ray photon is

captured by the nucleus of the atom with the ejection of a particle from the

nucleus when all the energy of the x-ray is given to the nucleus (absorption).

This process may be neglected for the energies of x-rays used in diagnostic

radiography.

Thomson scattering (R) (also known as Rayleigh, coherent, classical,

elastic scattering) occurs when the x-ray photon interacts with the whole atom

so that the photon is scattered with no change in internal energy to the

scattering atom, nor to the x-ray photon

X-ray dose concept and reduction measure

Hand X-ray of Mrs. Roentgen, spouse of Wilhelm Conrad Roentgen, the German

physicist who accidentally discovered unknown rays on November 8, 1895 and called

them... "X-rays".

This tutorial is intended to familiarize you with the dose concept. The purpose is to

give you a quick overview of the whole topic of x-ray dose through a simple

explanation of the technology and to expose the different techniques available to

reduce dose.

The Dose Concept

When an X-ray tube is in

operation, so-called X-ray

beams, a type of radiation,

are released. Using these

beams, the technician can

create images of whatever is

being examined. This

radiation penetrates objects

and human bodies, passes

through them, and is

weakened in the process.

In simple terms, this

weakening is equivalent to a

reduction in the number of

individual radioactive

particles. A statement

concerning the amount of

radiation, which is

measured at a site, produces the concept of "dose".

Because not all the radiation particles generated during an X-ray are used to produce

the resulting images, and because radiation can cause damage to the human body, we

Fig. 1: Determining Dose Parameters

try to achieve the greatest possible effect, that is the best possible image with the

smallest possible dose of radiation.

In general, the concept of "dose" can mean different things according to the

circumstance, for example according to the site where the dose is measured. For this

reason, the dose concepts most commonly used in radiology will be explained here.

Dose Parameters

Incident dose

The incident dose is the dose measured in the middle of a

radiation field on the surface of a body or a phantom.

However, it is only measured at this point if there is no body

in the path of the x-ray beam. Thus, there is no scatter

radiation from the body during this measurement. When

radiation strikes a substance, there is always a certain

scattering of radioactive particles. This is comparable to light

striking a glass surface; a certain portion of the light is

always reflected.

The unit used to measure the incident dose is joules per

kilogram, and is known as "Gray" where 1 Gray (Gy) = 1

J/kg. The former unit used to measure the incident dose was

the "Rad," and using this unit, 1 Rad (rd) = 0.01 Gy, or 1 Gy = 100 rd. But because

today's doses are generally very small, they are usually described using the unit "uGy",

that is, 0.000001 Gy.

Incident dose = the dose measured on the intended surface of the patient, but

without the presence of the patient

The System International unit (SI unit) used to measure the incident dose is the Gray,

where 1 Gy = 1 J/kg

Surface dose

The surface dose is measured with the body in the path of the beam.

Because of the scattered radiation that results on the surface and in the

depths of the body, the surface dose differs from the incident dose by

including the amount of scattered radiation.

Thus we can say:

Surface dose = incident dose + scattered radiation from the body

The SI unit used to measure the surface dose is the Gray (Gy)

Exit dose

The exit dose serves in the evaluation of the X-ray image. It is

measured in the radiation field in immediate proximity to the

surface of the body where the beams exit from the body. On the

basis of the exit dose and the surface dose, we can calculate how

much radiation must have remained in the patient's body.

Radiation in the body = surface dose - exit dose The SI unit used to measure the exit dose is the Gray (Gy)

Image receptor dose

The image Receptor dose is measured at the film cassette, X-ray

system's image intensifier assembly or Digital Detector. The

image receptor dose is generally smaller than the exit dose,

because the radiation weakens before it reaches the image receptor,

for example by encountering objects behind the patient's body

such as the radiation protection grid, anti-scatter grid or the table.

Image receptor dose <= exit dose

The SI unit used to measure the image receptor dose is the Gray (Gy)

Dose rate at image receptor

In order to measure a dose, the beam must operate for a certain period of time. The

dose rate therefore represents the measured dose for the amount of time required to

complete the dose measurement. If the image receptor dose is measured in the process,

then the dose rate is the image receptor dose rate. If the dose is measured at a different

site, then the dose rate is determined using one of the previously mentioned dose

parameters.

measured dose

Dose rate = -------------------------

required time

The SI unit used to measure the dose rate is Gray per second: (Gy/s) or (mGy/s)

Dose-area product

The dose-area product is a measurement of the amount of

radiation that the patient absorbs. It is usually measured behind

the multi-leaf collimator, that is, on the side of the patient where

the radiation enters the body, by attaching a measuring device in

front of the X-ray tube and passing a beam through it. The dose-

area product is independent of the distance between the X-ray

tube and the measuring device because the further away from

the X-ray tube this measurement is taken, the more the size of

the device increases, and the dose itself decreases (see diagram).

The dose to the patient can be calculated from the dose-area

product, the size of the measuring device, and the distance to the

X-ray tube and the patient.

Dose-area product = dose * surface area of the measuring device The SI unit used to measure the dose-area product is the Gray * centimeter2 (Gy*cm2)

Fig. 2:

Dose-

area

produ

ct

The

dose-

area

produ

ct at

50 cm

from

X-ray

tube is just as great as dose-area for 100 cm or 200 cm, because the size of the

measuring device increases with greater distance to the X-ray tube. But the dose itself

decreases with greater distance to the tube. Thus the dose-area product is the same at

each position if the size of the measuring device enables it to detect all of the radiation.

Body dose and effective dose

The body dose is the comprehensive concept for the organ or

partial-body dose equivalent and the effective dose. In the

practical application of radiation protection, however, local and

individual doses are monitored, because body doses cannot be

measured directly. The Radiation Protection Regulations

therefore use the concept of effective dose, in which all the

individual doses to the irradiated organs or parts of the body are

multiplied by a factor and then added together. The resulting

value may not exceed the dose limit for the effective dose that a

patient is allowed to receive.

Body dose = sum of all organ or partial-body doses

Effective dose <= patient dose limit

The SI unit used to measure the body dose and the effective dose is the sievert, where

1 sievert = 1 Sv = 1 Joule/kilogram = 1 Gray

Legal provisions

In many countries or states, by means of rules, guidelines or regulations, lawmakers

have contributed to improving radiation protection for patients and medical personnel;

after all, medical exposure to radiation is the single largest source of radiation

exposure among the general population. On an international level, guidelines are laid

down by the International Commission on Radiological Protection (ICRP). Many of

the rules, guidelines or regulations are governed by the ALARA concept (As Low As

Reasonably Achievable), meaning the production of a diagnostically relevant image

at minimum possible dose.

Special cases

Pediatrics

Because children have a greater sensitivity to radiation than adults, special conditions

apply to pediatric radiology. In particular, an attempt should be made to avoid X-ray

examinations altogether and to use alternative procedures instead, such as nuclear spin

tomography or sonography. Erroneous X-ray exams should be avoided and the dose

measurement should be taken with special pediatric measuring devices. Special X-ray

intensifying screens should be used as well to reduce additional dose. And because

children have a lower body thickness, the operator generally does not use an anti-

scatter grid, which is used when adults are X-rayed. Of additional benefit is a more

precise collimation of the area through which the beam will pass; this also reduces the

dose. Beam filtration is performed with a pediatric filter consisting of 0.1 mm copper

and 1 mm aluminum. It is especially important to use a gonad shield and to time a

child's inhalations exactly when making thoracic X-ray images.

Cardiology

Cardiology is another special case. The generator output must reach at least 100 kW,

and there must be an additional filter of 0.1 mm copper available for fluoroscopy as

well as an appropriate system of collimators for the radiation area. This should

include an iris diaphragm, and rectangular and semi-transparent collimators.

Calculation of the dose-area product during application has been prescribed by some

laws.

General dose reduction measures

Inverse square law

A bundle of X-rays corresponds to the shape of a cone, with the tube at its tip. The

intensity or dose of the radiation emitted from the source of the X-ray beam

diminishes with the square of its distance from the source. If you double the distance

x, the dose changes by a factor of 1/(2²), and if you triple it, the dose changes by a

factor of 1/(3²).

In general, the dose amounts to 1/x². Therefore, if you double the film-to-target

distance, you will need four times as much radiation to achieve the same image

blackening. If you did not change the patient's position, this would lead to radiation

stress in the patient; thus, increasing the distance between X-ray tube and patient

helps to reduce the dose.

Collimation at film

Collimation at the site of the film cassette does not result in any dose reduction,

because the radiation is not collimated to the appropriate film format until it has

passed through the patient. It merely serves to improve image quality by reducing

scattered radiation and thereby improving contrast.

Collimation at target

Collimation at the target brings about a genuine dose reduction and also produces

better image quality. Collimation is performed using cones and collimators (multi-leaf

collimators or iris diaphragms) that are attached directly in front of the X-ray tube.

Collimation at the target is the most effective radiation protection for the patient and

personnel, because it narrows the area that the radiation can strike.

Compression

Because radiation scatters in a body exposed to X-rays, compression of the body is

another way to reduce the radiation dose. Scattered radiation also produces an

undesirable reduction in contrast in the X-ray image. With compression, the thickness

of the body is reduced, and so a lower dose is absorbed by the body. Additionally,

compression ensures that less scattered radiation occurs.

Anti scatter grid

Fig. 3: Inverse square law

The anti-scatter grid is located

between the patient and the image

intensifier, or Cassette or Digital

Detector. It is the most effective

method of reducing scattered

radiation. The grid absorbs a

portion of the scattered radiation

in its lead plates.

This absorbed dose therefore does

not reach the image intensifier,

Cassette or Digital Detector, even

though it has already passed

through the patient. Thus, the use

of an anti-scatter grid leads to an

increase in the dose, because the amount of radiation that reaches the image intensifier,

cassette or Digital Detector, is not reduced until it has passed through the patient: if

the anti-scatter grid is used, the patient must be exposed to a higher dose of radiation

in order for the minimum dose to reach the Image intensifier, Cassette or Digital

Detector.

We can differentiate between individual anti-scatter grids using their grid ratios. This

is the relationship between the height of the plates to their distance from each other.

The greater the grid ratio, the greater the grid's effect. Thus, the required dose

increases with the grid ratio. The typical grid ratio is 8:1 for Radiography and 5:1 for

Mammography.

Kilovolt adjustment

The adjustment of the kilovolt values at the operating console also has an important

effect on the dose, because if a high kilovolt setting is chosen, the radiation is

"harder," that is, richer in energy and more able to pass through the body. High

kilovoltage and strong filtration are therefore similar in their dose reduction effects,

except that image contrast decreases with high kilovoltage.

Fig. 4: Cross section of an anti scatter grid

Fig. 5: Rhodium and Molybdenum energy

For Mammography, the traditional X-ray

tube target material is molybdenum, but

some equipment feature an additional tube

target material, Rhodium or Tungsten, in order to slightly harden the X-ray beam to

better penetrate dense breast without compromising image quality or contrast.

Radiation filtration / hardening

The quality of the X-rays also plays a great role in the size of the administered dose.

X-ray radiation normally has so-called "hard" and "soft" particles, that is, particles

with a lot of energy and particles with little energy. Hard particles are better for the

patient, because they pass through the body. Soft particles, by contrast, get caught

inside the body because they are too weak to pass through and out of it. Therefore, it

is primarily soft radiation that creates unnecessary exposure to the patient. For this

reason, copper and aluminum (Molybdenum and Rhodium in the case of

Mammography) are used as filters in front of the X-ray tube. The soft radiation is

caught in the filter plates, and the remaining radiation emerging from the filter is

"harder." This additional filtration can also reduce the dose to the patient without

diminishing image quality, because in any case only the "hard" rays reach the image

intensifier, film cassette, or Digital Detector.

Because the GE Senographe DMR+

and 2000D (Mammography Systems)

feature a double Molybdenum /

Rhodium X-ray tube tracks as well as

two different filters, they provide a

good example of the impact of

different X-ray target/filtration

materials on dose (fig. 6).

Film/screen combinations

Choosing the right film/screen combination can greatly influence the required dose. In

general, the dose is a function of the sensitivity of the combination. This sensitivity is

spectrums for GE Mammography X-ray

tube.

Fig. 6: Target/filtration materials impact on

dose

the quotient of 1000 uGy and the required dose in uGy.

1000 uGy

Sensitivity = -------------------------

Required dose in uGy

or:

1000 uGy

Required dose in uGy = -------------------

Sensitivity

For example, a combination with a sensitivity of 400 requires 2.5 uGy in dose. The

sensitivity is greatly dependent on the intensifying screen that is used, because the

screen is the principal component in image blackening. We differentiate mainly

between screens using traditional fluorescent materials such as calcium tungstate, and

screens made of so-called rare earths. These rare earths intensify better, which means

they transform more X-ray beams into light. They can therefore reduce the dose by up

to 50%, because the operator can select a lower dose and still get the same image

quality that would be attainable using traditional screens. These screens are stipulated

in pediatrics, for example.

Image intensifier input screen

At the input screen of the image intensifier there is a situation similar to that of the

screen-film combinations. The input fluorescent screen substantially determines the

intensification, that is, the transformation of the X-rays into light. In conjunction with

the X-ray image intensifier, the age of the X-ray system plays a role, because the

properties of intensification decrease considerably with age. In addition, the radiation

field adapts automatically to the format of the image intensifier, which also lowers the

dose, because only small portions of the patient are irradiated instead of the patient's

entire body. The choice of a small input screen causes collimation, and this too leads

to a reduction in the dose.

Automatic exposure timing

The automatic exposure timer or Automatic Exposure Control (AEC) measures the

dose of radiation that strikes the X-ray film behind the patient, and turns the X-ray

system off when the predetermined dose for that screen-film combination has been

reached. This assures that only the smallest required dose is administered. The

resulting images all show a uniform blackening, and the danger is reduced that the X-

ray examination might have to be repeated owing to an error in the image. In this way,

automatic exposure timing also indirectly reduces the dose.

Automatic dose rate adjustment

The dose rate is the dose over the total time in which the patient is exposed to

radiation. If the radiation exposure time to the body can be reduced, this leads to a

decrease in the total dose to the patient. By automatic dose rate adjustment, the

operator tries to reduce the time during which the dose rate is measured at the input of

the image intensifier and the kilovolt and milliamp values are, in turn, adjusted at the

generator. In the process, the dose rate should be kept as low as possible. Automatic

dose rate adjustment is comparable to automatic exposure timing for images made

using a screen-film combination.

Tabletop

The material from which the tabletop is constructed is also significant for the required

dose, because the tabletop is penetrated by the radiation and weakens it before it

reaches the image intensifier. Therefore, if at all possible the tabletop should not

contain any material that strongly weakens the radiation or absorbs it well, such as

lead or metals in general. Carbon fiber has proven to be the best material for X-ray

system tabletops because its radiation absorption is minimal and the tabletop can take

a great amount of stress; today a tabletop is expected to be able to support a patient

weighing 120 - 150 kg (up to 330 lbs).

Special dose reduction measures Low-dose fluoroscopy

Fluoroscopy using a reduced dose has become possible primarily through digital

technology. Principally, parts of the body with low levels of spontaneous movement

are well suited to this method of examination. A few digital fluoroscopy procedures

will be described in the following paragraphs.

Pulsed fluoroscopy

In pulsed fluoroscopy, X-rays are no longer delivered continuously; they are delivered

in pulses that follow in rapid succession. This reduces the amount of time during

which radiation is released. The resulting radiation-free gaps in the imaging process

are filled with the last stored digital image until a new and more current image is

available. The short X-ray pulses mean that the dose is significantly reduced;

additionally, image definition is increased. Pulsing can take place either by using a

pulse control at the X-ray generator, or with a grid-controlled X-ray tube; however,

the grid control leads to a lower level of radiation exposure.

Fig. 7a: Pulse Fluoroscopy The illustration shows the advantages of grid control of the X-ray

Fig. 7.a Top, we

can clearly see

the exact pulses through the grid control; moreover, they allow rapid switching times.

Fig. 7a Bottom in contrast, we can see that during pulsing controlled at the generator,

the kilovolt values move more slowly toward the correct value and away from it again,

resulting in the patient receiving an unnecessary dose of pulses with a low kilovolt

value. Low kilovolt values contribute to radiation exposure but do not result in usable

images.

Grid control can itself be

subdivided into pulse

frequency control and

pulse width control (Fig

7b). The frequency of

pulse frequency control

can be varied for example,

12 b/s or 3 b/s, and it

controls the X-ray tube

continuously. Pulse width

control, on the other hand,

changes the duration of the

individual pulses while at a constant frequency, for example, 25 b/s.

Image integration

Image integration means adding together two or more individual images to create a

single image. The dose rate can either be kept the same, resulting in images that are

clearer because there is less noise in the image, or the dose can be reduced. This is

accomplished by reducing the dose rate and adding individual images to each other

until the same image quality is achieved without image integration, but with a

reduction in total dose. Combining individual images does result, however, in fewer

finished images for viewing at the monitor. The gaps that occur with this method are

filled by outputting the same image twice in a row, similar to the method used for

pulsed fluoroscopy. The reduction in dose with this method is approximately 50%. A

disadvantage of the method is the stroboscopic impression that can arise with fast

moving objects.

Digital filtering / SMART Fluoro

With digital filtering, also known as recursive filtering, a fluoroscopy image is mixed

or overlaid with one or more previously stored images. The proportion of the previous

images is smaller the farther back or longer ago that they were acquired. On the whole,

there is flexibility in choosing the proportional mixture of images, and the process

represents a compromise between dose reduction and the lag effects that result when

mixing images. However, the dose rate can be significantly lowered since the images

that are produced are only new, or newly made, to a certain degree; that is, combining

tube (top) compared to control at the generator (bottom).

Fig. 7b: Pulse Fluoroscopy

The Diagram shows the constant pulse width of the

selected frequency during pulse frequency control (top),

and the constant frequency with variable pulse width

using pulse width control (bottom).

individual images means that less overall dose is required. The image mixture

proportion can be monitored using a motion detector that lowers the image mixture

proportion in the event of a strong shift in the gray scale values in the image, for

example when there is movement, so that the output predominantly reflects the

current gray scale value.

Last image hold

Last image hold means that the last image obtained during a fluoroscopy is stored

until a new image is produced. The physician can then study the image without

further radiation exposure. This can lead to a reduction in radiation exposure since the

total fluoroscopy time is reduced, and with it, the total dose.

Frame grabbing

Frame grabbing means that the physician can "grab" or extract and view a chosen

image from a fluoroscopic series without the necessity of additional radiation.

Additionally, spot film radiography can be reduced because the physician can use the

"frozen" images from the fluoroscopic series. This dose reduction is particularly well

suited to pediatrics.

Roadmapping

Roadmapping is the overlaying of two images. A stored image is superimposed upon

a current fluoroscopic image, or a current image can be copied for storage and later

used in roadmapping. This is useful primarily in viewing blood vessels, because an

existing image of a blood vessel filled with contrast medium can be superimposed on

a catheter image made during fluoroscopy. This can save time and contrast medium,

and reduce the radiation dose.

Digital Fluoro imaging techniques

Due to digital correction functions and the superior quality of modern X-ray image

intensifiers, it is possible to produce spot films from previously captured fluoroscopy

images instead of making new images with screen-film systems. Thus, the additional

dose that would be needed for new spot film radiography can be completely

eliminated, which means a significant reduction in dose. The somewhat lower quality

of the fluoroscopy-generated images, which stems from the lower dose used for

fluoroscopy, is usually accepted as a reasonable compromise.

Dose levels

The subdivision of the dose into individual levels permits finer gradation. With this

practice, the minimum dose for optimum image quality can be selected for every kind

of examination. In addition, the standard examination protocols can be individually

adapted to the patient.

Virtual collimation

Virtual collimation is a term used to describe the possibility of positioning the

collimators via a display or at the monitor. Pre-setting the collimators this way does

not require an X-ray beam, and thus reduces the time that the patient is exposed to

radiation and consequently, the entire administered dose.

Solid State detector

Another possibility for dose reduction is provided by the use of an electronic flat bed

detector, also known as a solid state detector. These silicon-base detectors have a

higher degree of effectiveness than traditional detectors, which is expressed in

Detective Quantum Efficiency (DQE). The patient dose is in direct proportion to this

quantum efficiency:

Image Quality

Patient Dose proportional to -------------------------

Detective Quantum

Efficiency (DQE)

On the basis of this equation, we can see that the greater the Detective Quantum

Efficiency, the smaller the dose for the patient, yet with the same image quality. An

electronic flat bed detector therefore means that a larger amount of the released

radiation is actually used, so that from the outset a smaller dose of radiation is needed

to produce a comparable image.

Also, because Solid State detectors feature a high dynamic range, they can

accommodate for less X-ray and compensate for the lack of film blackening through

appropriate Brightness and Contrast adjustment techniques. For some Solid State

detectors, the Image Quality does not suffer from a higher energy X-ray beam, thus

contributing to a decrease of the overall patient dose.

GE X-ray equipment and dose reduction measures

Revolution XQ/i and XR/d Digital X-ray Systems

The digital radiography systems from GE are designed to

meet your clinical requirements today and to address the

trends that will make digital x-ray a logical imperative over

the next decade.

The Revolution XQ/i is designed to improve clinical

effectiveness and productivity of Chest Exams

The Revolution XR/d includes a four-way float-top

elevating table.

The Revolution XQ/i and XR/d feature a high DQE that can contribute to reduction of

dose.

Senographe 2000D and DMR+

Senographe 2000D is the new Digital Mammography System

from GE Medical Systems. It is a complete, modular system

that eliminates the need for film cassettes and takes full

advantage of digital technology from on-screen image display

to Networking, Filming and Archiving.

The Senographe DMR+ employs a unique, patented bi-metal

mammography tube with a Rhodium track for superior

imaging of the most challenging breast tissues.

The Senographe 2000D feature a Solid State Detector with High DQE in addition to

the dose reduction measures already incorporated into the Senographe DMR+:

Additional Rhodium target X-ray tube

Rhodium filter

X-Ray Image Intensifier

The x-ray image intensifier has been used for almost fifty years to produce sequences

of x-ray images. Its origins lie in low light level imaging, for example, night vision

devices, to which an x-ray intensifying screen has been added.

The construction, mode of operation and performance characteristics of x-ray image

intensifiers are considered on this page, under the following headings:

Construction and Mode of Operation

The x-ray image intensifier (XII) is generally a cylindrically-shaped device containing

a number of components housed in a vacuum. Figure 4.1

Fig 4.1

shows a cross-section through this cylinder. X-rays emerging from the patient enter at

the input window and strike the input phosphor. The input phosphor scintillates and

light photons strike the photocathode, which emits electrons. These electrons are

accelerated and focussed by the electron optics onto the output phosphor which emits

light. This light provides an image of the x-ray pattern that emerged from the patient

which has a substantially greater intensity than when an intensifying screen is used on

its own. This description of its operation is summarised in figure 4.2 .

Fig 4.2

The major components inside the XII include:

Input Window

The input window in older XIIs was made from glass and their performance

suffered from x-ray scattering and absorption effects in this material. This

limitation has been overcome in modern devices by using a relatively thin

sheet (e.g. 0.25 - 0.5 mm) of aluminium or titanium where good strength is

achieved for containing the vacuum with minimal x-ray attenuation.

Input Phosphor

The input phosphor is made from CsI, doped with Na, which is deposited on

an aluminium substrate. The CsI:Na is grown in a structure of monocrystalline

needles, each about 0.005 mm in diameter and up to 0.5 mm long. The

aluminium substrate is about 0.5 mm thick (figure 4.3 - note that dimensions

in the figure are not to scale). The input phosphor is typically 15 to 40 cm in

diameter, depending on the XII.

Fig 4.3

Both Cs and I are good absorbers at diagnostic x-ray energies having K-edges

at 36 and 33 keV, respectively. The CsI:Na phosphor produces a blue light

when x-rays are absorbed and this light is guided along the needles in a fibre-

optic fashion (i.e. without much lateral spread) to the photcathode.

A photograph of a glass-envelope XII which has been cut in half to expose the

inner side of the input components is shown in figure 4.4.

Photocathode

An intermediate layer (less than 0.001 mm thick) is evaporated onto the inner

surface of the CsI:Na phosphor and a photcathode (about 2 nm thick) is

deposited on this layer (figure 4.3).

The intermediate layer (e.g. indium oxide) has a high optical transmission and

is used to chemically isolate the phosphor and photocathode materials. The

photocathode typically consists of an alloy of antimony and caesium (e.g.

SbCs3).

Light photons emitted by the input phosphor are absorbed via the

phototelectric effect in the photocathode to release photoelectrons.

Vacuum & Electron Optics

The vacuum is required so that the electrons can travel unimpeded - as in the

case of the x-ray tube. A voltage of 25 to 35 kV is used to accelerate the

electrons and the electon optics is used for focussing them onto the output

phosphor. A current of about 10-8

to 10-7

A results and it is the acceleration

and focussing of these electrons which gives rise to the image intensification.

Note that a cross-over point exists so that the image at the output phosphor is

inverted relative to that at the input phosphor. Note also that the input

phosphor and photocathode are in fact curved ( i.e. not perfectly straight as

shown in figure 4.1) so as to equalise the electron path lengths and hence

minimise image distortion.

Image magnification can be achieved by varying the voltages on the electrodes

of the electron optics, so that a 38 cm XII can also be used to image field sizes

of 26 cm and 17 cm, for instance. Three discrete field sizes are typical of

many systems although XIIs with a continuous zoom feature are also available.

Image brightness decreases as the field size is reduced when the input

exposure rate is maintained constant.

Most XIIs also feature mechanisms for establishing and maintaining the

vacuum, but this aspect of their construction is beyond the scope of the

treatment here.

Output Phosphor

The output phosphor is made from ZnCdS: Ag (e.g. a P20 phosphor) deposited

on the ouput window (figure 4.5 - note that dimensions, once again, are not to

scale).

This phosphor emits a green light when it absorbs the accelerated electrons,

and is typically about 0.005 mm thick and 25 to 35 mm in diameter.

In addition, a thin aluminium film is placed on the inner surface of the

phosphor, which serves both as the anode and to reflect light back towards the

output window - so as to increase the output luminance and to prevent these

light photons exciting the photocathode.

Output Window

A number of designs of output window exist and include a glass window (e.g.

15 mm thick) with external anti-reflection layers, a tinted glass window and a

fibre-optic window - the objective of these designs being to minimise light

diffusion and reflections.

The resulting image is fed to an optical system to be viewed by a cine-camera,

photographic camera, video camera or combinations of these cameras.

Orthochromatic film is needed for the film-based cameras.

In summary, consider the fate of a 50 keV

x-ray photon which is totally absorbed in

the input phosphor:

o The absorption will result in about 2,000

light photons, and about half of these might

reach the photocathode.

o If the efficiency of the photocathode is

15%, then about 150 electrons will be

released.

o If the acceleration voltage is 25 kV, the

efficiency of the electron optics is 90% and

each 25 keV electron releases 2,000 light

photons in the output phosphor, then about

270,000 light photons will result.

o Finally, if 70% of these are transmitted

through the output window, the outcome is

a light pulse of about 200,000 photons

produced following the absorption of one

50 keV x-ray.

The XII envelope is made from glass or non-magnetic stainless steel, and the input

window is welded to this envelope. The assembly is housed inside a metal container

which contains lead, for radiation shielding, and mu-metal, to shield the electron

optics from external magnetic fields. The input window is typically protected by an

aluminium faceplate (e.g. 0.5 mm thick) which also serves as a safety device in case

of implosion of the XII. Many systems also feature a scatter-reduction grid mounted

at the faceplate. A 15 cm XII assembly is shown in figure 4.6, with the faceplate at

the top of the photograph, and the optical system and a video camera towards the

bottom.

Performance Characteristics

Brightness Gain

The gain in image brightness results from the combined effects of image

minification and the acceleration of the electrons:

o Minification Gain

This results because electons from a relatively large photocathode are

focussed down to the smaller area of the output phosphor which gives

rise to an increase in the number of electrons/mm2. The gain is given

by the ratio of the areas of the input and output phosphors and can be

expressed as:

Thus.....

for input phosphors with diameters between 15 and

40 cm

and an output phosphor of 2.5 cm diameter,

the minification gain is between 36 and 256.

o Flux Gain

This results from the acceleration given to the electrons as they are

attracted from the photocathode to the output phosphor. It is dependent

on the applied voltage and is typically between 50 and 100.

Brightness Gain

The overall brightness gain is the product of the minification gain and

the flux gain, i.e.

.

Brightness Gain = (Minification Gain) x (Flux Gain) .

Thus, when:

Minification Gain = 100 and

Flux Gain = 50

then the Brightness Gain = 5,000

Brightness Gains to more than 10,000 are

achievable

Conversion Factor

The brightness gain is not easily measured and serves simply to illustrate the

performance of an XII. A more readily measured parameter is the conversion

factor, which relates what the XII delivers (i.e. luminance) relative to the input

(i.e. radiation exposure), and is useful for comparing the performance of XIIs

as well as that of a given XII over time. The output luminance is measured

using a photometer, the radiation exposure with a ionization chamber and the

conversion factor is expressed as:

This factor is typically 7.5 - 15 Cd m-2

/µGy s-1

and higher.

Note that the resulting image is relatively dim. For example the luminance of a

standard domestic light bulb is about 106 Cd m

-2. The green output image

therefore needs a darkened room and dark-adapted eyes for direct viewing, or

a sensitive video camera for remote viewing.

Contrast Ratio

This parameter expresses the broad-area, high contrast performance of an XII.

Various scattering effects inside the XII result in a radio-opaque object not

being completely opaque in the image. The contrast ratio can be measured by

imaging a lead disk and expressing the luminance of its image relative to that

of an open field image. For standardization purposes, the size of the disk is

typically 10% of the field size and it is placed centrally in the field of view. It

is expressed as follows:

Typical values are 20:1 to 30:1 or greater.

The contrast is affected by factors which include:

o X-ray scattering in the input window

o X-ray scattering in the input phosphor

o Light scattering in the input phsophor

o Electron scattering in the electron optics

o Light scattering in the output phosphor

o Light scattering in the output window

and these scattering effects are collectively referred to as veiling glare.

The output phosphor has generally been regarded as the major source

of veiling glare although recent work has indicated that the input

components may also contribute significantly.

Limiting Spatial Resolution

This parameter can be assessed using a Pb bar test pattern by determining the

highest spatial frequency - in line pairs per mm (lp/mm) - that can be resolved.

Images of such a test pattern are shown in figure 4.7, where a 23 cm XII is

operated in a 23 cm (left), a 15 cm (middle) and an 11 cm (right) mode. The

parameter is generally expressed for the centre of the field of view, since it

decreases towards the image periphery depending on the quality of the

electron optics.

The performance is dependent on the field size and the type of imaging

camera used, and the table below shows some typical results. Note that each of

the cameras degrades resolution to some extent and that the XII itself has a

lower resolution than screen/film devices.

Field

Size

(cm)

Output

Phosphor

100/105

mm

Camera

35 mm

Camera

Conventional

Video System

15 - 18 5 lp/mm 4.2 lp/mm 2.5 lp/mm 1.5 - 1.3 lp/mm

23 - 25 4.2 lp/mm 3.7 lp/mm 2.2 lp/mm 1.0 - 0.9 lp/mm

Spatial Non-Uniformity

XII images of a uniform object are generally brighter in the centre than in the

periphery due to an unequal brightness gain in different regions of the field of

view.

This effect is also called vignetting and is illustrated in figure 4.8. The image

is of a uniform object acquired with an XII coupled to a video camera, and the

graph on the right shows the brightness profile for a horizontal line through

the centre of the image. The vignetting effect is quite apparent.

Note that this parameter is not widely assessed for XII systems - in contrast to

nuclear medicine where image uniformity of gamma cameras is rigorously

controlled - and a standardized measurement technique is not in widespread

use. Note also that the image data in the figure reflects the combined effects of

the spatial uniformity of the detected radiation beam, the coupling optics and

the video camera, and not solely the XII.

Spatial Distortion

The final performance characteristic to be considered is spatial distortion. All

XIIs suffer from this effect, where images do not faithfully reproduce the

spatial relationships in an object because of unequal magnification in different

regions of the field of view.

The effect is illustrated in figure 4.9, where images of a regular wire matrix

acquired with a modern (on the right) and an older XII are shown. The

distortion is typically 'pincushion' in nature - as readily seen in the image on

the left. Notice, for example, that straight lines are reasonably straight in the

centre of this image and change to curves towards the peripheral regions.

Notice, also, that image areas measured in the centre of the field will be less

than those measured in the periphery.

One approach to assessing this characteristic is to determine the integral

distortion, which is expressed as follows:

where:

o D1: diagonal length of the central square in the image of the matrix

o D2: diagonal length of the largest square in the image

o n: a factor to account for the relative sizes of these squares in the object.

The distortion expressed using this approach is 8.5% for the image on the right

in figure 4.9, and 3.5% for the image on the right.

Film-screen Speeds

The sensitivity of a film-screen combination depends on the film, the screen, the film

processing, and the beam quality, i.e. the spectrum of the X-rays exposing the film

screen combination. This explains immediately, why the sensitometry of a film-screen

combination with X-rays is a lot more complex than the sensitometry of a film with

light, and therefore is hardly ever done outside the manufacturer's laboratory:

1. The film-screen combination has to be exposed with a standardized spectrum. This

requires the use of a specified high voltage value, a specified high voltage waveform

(usually DC), a specified target composition, a specified filtration, all resulting in a

specified half-value layer.

2. While the film-screen combination has to be exposed with different dose values,

the operating parameters of the X-ray source (tube voltage, tube current, and exposure

time) must not be changed, as this is the only way to avoid measurement errors due to

spectral changes and due to the reciprocity law failure. Therefore, the dose can only

be varied by changing the distance between source and film-screen combination.

3. The film has to be processed under standardized conditions.

The speed of a film-screen combination is stated as the inverse of the dose (in Gy)

needed to obtain a film density of one above base plus fog, multiplied by 1000 Gy:

1000 Gy

SPEED = -------------------------

Dose for D = 1+Base+Fog

The speed is the quotient of two dose values, it does not have a dimension or unit

name attached to it. As the speed is inverse proportional to the dose requirements of a

film-screen combination, twice the speed is equivalent to half the dose and vice versa.

With this definition, the standard or universal film-screen combinations with calcium

tungstate phosphor used to have a speed of 100. With the modern rare-earth systems,

the speed of the standard screen is usually 200, i.e. the film-screen combination for

universal application requires 5 Gy (approximately 0.5 mR) for a film density of

one plus base plus fog. The speed values of the high resolution ("detail" or "fine")

resp. the high sensitivity ("high speed") film-screen combinations of one and the same

product line differ from the speed (and thus, dose requirement) of the standard

combination by a factor of two in either direction. Thus, a rare-earth "detail" film-

screen combination has a speed of 100, and a rare-earth "high speed" film-screen

combination has a speed of 400. These are typical values, but for special applications

screens with lower and higher speeds are available.

Relay

It is often desirable or essential to isolate one circuit electrically from another, while

still allowing the first circuit to control the second. For example, if you wanted to

control a high-voltage circuit from a control desk, you would not want to connect it

directly to control panel in case something went wrong and the high voltage became

connected to the control desk.

One simple method of providing electrical isolation between two circuits is to place a

relay between them, as shown in the circuit diagram of figure 1. A relay consists of a

coil which may be energised by the low-voltage circuit and one or more sets of switch

contacts which may be connected to the high-voltage circuit.

How Relays Work

In figure 2a the relay is off. The metal arm is at its rest position and so there is contact

between the Normally Closed (N.C.) switch contact and the 'common' switch contact.

If a current is passed through the coil, the resulting magnetic field attracts the metal

arm and there is now contact between the Normally Open (N.O.) switch contact and

the common switch contact, as shown in figure 2b.

Advantages of Relays The complete electrical isolation improves safety by ensuring that high voltages and

currents cannot appear where they should not be.

Relays come in all shapes and sizes for different applications and they have various

switch contact configurations. Double Pole Double Throw (DPDT) relays are

common and even 4-pole types are available. You can therefore control several

circuits with one relay or use one relay to control the direction of a motor.

It is easy to tell when a relay is operating - you can hear a click as the relay switches

on and off and you can sometimes see the contacts moving.

Disadvantages of Relays

Being mechanical though, relays do have some disadvantages over other methods of

electrical isolation:

Their parts can wear out as the switch contacts become dirty - high voltages and

currents cause sparks between the contacts.

They cannot be switched on and off at high speeds because they have a slow response

and the switch contacts will rapidly wear out due to the sparking.

Their coils need a fairly high current to energise, which means some micro-electronic

circuits can't drive them directly without additional circuitry.

The back-emf created when the relay coil switches off can damage the components

that are driving the coil. To avoid this, a diode can be placed across the relay coil.

X-Ray beam modification in general radiography

Filtration

1) Filter: piece of metal (typically aluminum) located between the x-ray tube and the collimator box

and in the path of the primary beam

2) Purpose: to remove non-diagnostic, low-energy photons from the primary beam which in turn

reduces skin dose to the pt

3) Filters will cause partial absorption/attenuation of the x-ray beam

a) Attenuation/absorption: reduction in the total number of x-ray photons remaining in the beam

after passing through a given thickness of material

4) Inherent vs added filtration

a) Inherent: caused by glass, oil, tube housing, port or window

i) Roughly 0.5 mm Al equivalent

b) Added: caused by collimator – thin sheets of ~1.0 mm Al

c) NCRP recommendations

i) 2.5 mm Al equivalence filtration for tubes operating above 70 kVp

5) Filtration will reduce exposure rate

6) Filtration affects beam quality/energy/penetrating ability

a) Average beam energy/penetrability

i) Depends on:

(1) kVp

(2) Amount of total filtration in the beam

ii) kVp also determines the minimum wavelength of the beam

iii) Filtration determines the maximum wavelength of the beam

iv) Increasing either kVp or filtration will increase the average energy of the beam, allowing

it to be more penetrating and of higher quality

b) Beam quality is measured by its half-value layer (HVL)

i) HVL directly measures beam quality by determining actual penetrating ability

(1) Federal regulation states that at 80 kVp, the half-value layer must be 2.34 mm

equivalent

(2) Mammography has different standards because of the desire to keep the softer x-rays;

regulation call for an HVL of 40 mm equivalence at 30 kVp

ii) HVL: that thickness of a specified material (usually a metal) which reduces the exposure

rate to one-half its initial value

iii) The HVL principle is utilized when extending technique charts

c) Because filtration causes the beam to be more penetrating, increasing filtration:

i) Decreases density

ii) Decreases exposure rate

iii) Decreases contrast

7) Types of filters

a) Thoreau‘s

i) Compound filter used in therapy

ii) Compound materials include tin, copper, and aluminum

iii) 250 – 450 kVp

iv) The layering order of these metals (tin is closest to the tube, aluminum closest to the pt)

is important due to characteristic radiation; new x-rays formed in the first layer are

absorbed by the next layer; aluminum‘s characteristic radiation is absorbed in the air

(1) Characteristic radiation

(a) Incoming electron collides with an inner shell electron of the target material,

displacing that electron from it‘s shell

(b) An electron from a higher shell will drop down to fill the newly created space

(c) Energy given off is a characteristic x-ray, called characteristic because its

energy is characteristic of the target element and its involved electron shell

b) Compensating filter

i) May be made of metal or a plastic compound (EX: boomerang)

ii) Used where there is difficulty imaging body parts due to varying tissue thickness and

composition

iii) A wedge filter, shaped as it is named, allows for greater attenuation of the beam at its

thicker end

(1) Usually made of aluminum

iv) A trough filter, lower in the middle than at sides, is used in chest radiography to allow for

greater filtration over the lung tissue and less over the mediastinum

v) Computer radiography incorporates its own compensating filtration

Scatter Radiation

1) Factors affecting the amount of scatter:

a) Patient thickness

b) Tissue density

i) Total volume of body tissue = length x width x height

(1) Length is determined by the thickness of the part

(2) Height and width are determined by collimation

(3) Controlling tissue volume is done via tissue compression and/or adjusting field size

c) Field size

d) KVp

i) As kVp is increased, more energy is able to reach the film and so more scatter is

produced

Beam Limitation

1) Volume of tissue determines the amount of s/s radiation

a) Volume = thickness x area

2) Increased collimation means:

a) Decreased volume of tissue irradiated

b) Decreased s/s radiation

c) Decreased fog

d) Decreased density

e) Increased contrast

f) The effects of collimation are more evident with thick body parts and non-grid exposures

3) Beam limitation protects the patient from unnecessary radiation

4) Increasing beam limitation will decrease density, with all other factors constant

5) Beam limitation improves visibility of detail with technique compensation

a) Needed only with extreme increases in collimation, such as going from a collimation of

14x17 to collimation of 5x5

6) Beam limitation is the most effective method for limiting scatter

7) Beam limiting devices include

a) Aperture diaphragm

i) Essentially a metal disk with a hole in its center

ii) Major disadvantage is that the aperture diaphragm allows more penumbra and off-focus

radiation

b) Collimator

i) Comprised of 2 independently-acting sets of adjustable lead shutters

ii) A mirror angled 45° and light bulb are set up to indicate alignment of the central ray

(1) If the mirror‘s angulation is off, the collimator light will not be true to the actual

exposure field

iii) Adjustable shutters allow collimated shapes to match the shapes of cassettes

iv) Helps to limit penumbra

v) The collimator is the most effective of beam limiting devices

c) Cone

i) Disadvantage includes allowing a penumbra

ii) Cones are useful for headwork, L5-S1 spot, sunrise, and other small parts

8) Positive beam limitation (PBL): automatic collimation which automatically adjusts to the cassette

size

Grids

1) The purpose of the grid is to absorb scatter and increase image contrast

2) Grids are located between the patient and the film

3) Grids absorb scatter which has already been produced

4) Construction:

a) Thin lead strips alternate with interspacing material

b) Interspacing

i) Organic (carbon-based) interspacing absorbs moisture and can potentially warp (EX: fiber,

paper, cardboard, plastic)

ii) Inorganic interspacing is much more durable and absorbs more radiation (EX: aluminum

and the less-visible lead)

5) Types of grids

a) Linear

i) Comprised of one set of lead strips extending in parallel fashion in one direction

ii) Strips are aligned with the long axis of the grid or the long axis of the table

b) Crossed / cross-hatch

i) A second set of lead strips is set perpendicular to the first set

ii) This grids will not allow for the use of any tube angle

c) Parallel

i) Lead strips are set parallel to one another

ii) These grids allow cut-off along the edges at shorter SIDs

d) Focused

i) Grid strips are angled progressively as they move further from the grid center in order to

coincide with the shape of the beam

ii) Convergence line: imaginary line in space created by extending the edges of angled lead

strips until they meet

iii) Grid radius / focusing distance: distance from the convergence line to the grid

(1) A focal range will be given on the grid

iv) Stationary and bucky grids use linear focused grids

e) Rhombic

i) A type of crossed grid in which grid strips are angled with respect to one another

6) Grid characteristics

a) Grid ratio

i) Grid ratio = height of lead strips / distance between strips

ii) The grid ratio indicates how well the grid cleans up scatter

(1) Higher ratios mean higher absorption of scatter

iii) Higher grid ratio means greater need for precision when centering in order to avoid grid

cut-off (increased ratio means decreased latitude)

iv) As grid ratio increases, mAs will need to be increased to maintain density

v) As grid ratio increases, contrast will increase

b) Bucky factor / grid factor (bf)

i) The bucky factor defines the requirement for increasing exposure factors to maintain

density with the use of a grid

c) Grid frequency (gf)

i) Grid frequency indicates the number of lead strips in an inch or centimeter

ii) As frequency increases, the strips get thinner

iii) Grid frequencies most used in diagnostic radiography are 85 – 103 lines per inch

iv) Thinner strips are not as visible on images, but they are not as effective in cleaning up

scatter

v) If two grids have an equal ratio, the one with the fewer, and thus thicker, strips will be

the more efficient grid, although its gridlines will be more visible

d) Contrast improvement factor (gk)

i) GK = contrast with a grid / contrast without a grid

ii) Useful GK numbers range from 1.5 to 3.5

iii) As the grid factor increases, the contrast improvement factor increases

e) Grid selectivity (gΣ)

i) Grid selectivity = % of primary beam transmitted / % of scatter transmitted

ii) This number describes grid efficiency

iii) Grids absorb around 20% to 40% of the primary beam

7) Grid selection and use

a) Use a grid with body parts measuring 10 centimeters or more

b) Use a grid with kVp values over 60

Air Gap

1) Air gap is defined by a 6‖ to 10‖ OID

2) Air gap may be used in consideration of scatter reduction over use of a grid since the space

traversed by scatter radiation allows it to miss striking the image receptor

3) One disadvantage of air gap technique is magnification

Rectification

Rectifier

The purpose of a rectifier is to convert an AC waveform into a DC waveform. There

are two different rectification circuits, known as 'half-wave' and 'full-wave'

rectifiers. Both use components called diodes to convert AC into DC.

A diode is a device which only allows current to flow through it in one direction. In

this direction, the diode is said to be 'forward-biased' and the only effect on the signal

is that there will be a voltage loss of around 0.7V. In the opposite direction, the diode

is said to be 'reverse-biased' and no current will flow though it.

The Half-wave Rectifier

The half-wave rectifier is the simplest type of rectifier since it only uses one diode, as

shown in figure 1.

Figure 2 shows the AC input waveform to this circuit and the resulting output. As

you can see, when the AC input is positive, the diode is forward-biased and lets the

current through. When the AC input is negative, the diode is reverse-biased and the

diode does not let any current through, meaning the output is 0V. Because there is a

0.7V voltage loss across the diode, the peak output voltage will be 0.7V less than Vs.

While the output of the half-wave rectifier is DC (it is all positive), it would not be

suitable as a power supply for a circuit. Firstly, the output voltage continually varies

between 0V and Vs-0.7V, and secondly, for half the time there is no output at all.

The Full-wave Rectifier

The circuit in figure 3 addresses the second of these problems since at no time is the

output voltage 0V. This time four diodes are arranged so that both the positive and

negative parts of the AC waveform are converted to DC. The resulting waveform is

shown in figure 4.

DC TESLA COIL DESIGN

This page discusses the application of a DC supply to Tesla Coiling. This page covers

the simple resistive charging arrangement and contains a link to the more complex

DC resonant charging topology. The latter was used by Greg Leyh in his Electrum

coil design.

Although the title says "DC Tesla Coil Design", the Tesla Coil itself is still inherently

an AC device. The tank capacitor still sees a polarity reversal during ringdown,

regardless of the type of supply used to charge it. However, there are several benefits

to using a DC supply to charge the tank capacitor.

As with AC charging, a high voltage power supply is used to charge the tank

capacitor of the Tesla Coil. However the main difference is that the source of power is

a smooth DC supply, rather than an AC supply operating at the mains frequency. This

results in some significant differences in behaviour, most of which are advantageous.

(See advantages and disadvantages section later.) In particular the removal of the line

frequency from the charging circuit can allow the firing rate of the rotary spark gap to

be varied over a wide range without encountering any beating or surging problems.

The DC charging arrangement can be broken down into two separate stages:

1. The High Voltage DC supply. (HVDC)

2. The charging circuit.

Both of these stages will be described in detail

in the sections that follow…

HVDC SUPPLY

The job of the HVDC supply is to provide a constant high voltage output of fixed

polarity to the charging circuit that follows. A perfect supply would provide its rated

voltage with no ripple, and its output voltage would not drop when current is drawn

from it. In practice this ideal supply can rarely be realised, and a compromise must be

made. It is in the areas of ripple and regulation where this compromise is made.

There are many different ways to build a HVDC supply. Possible designs stretch from

simple single phase supplies to elaborate 3-phase arrangements. There is a trade-off

between simplicity of the design and the performance. Some of the more common

alternatives are shown below.

Single phase supply

One of the simplest arrangements for a HVDC supply is shown below:

This design uses a step-up transformer followed by a bridge rectifier and a smoothing

capacitor. The rectifier converts the high voltage AC from the transformer into pulsed

DC, and the smoothing capacitor acts like a reservoir and holds the peak voltage for

the time between peaks

The DC output voltage from this arrangement is equal to 1.41 times the RMS voltage

rating of the transformer. (i.e. It is equal to the peak output voltage from the

secondary winding of the transformer.)

Although this circuit is cheap and simple, it exhibits a few shortfalls. The output from

the rectifier pulses at twice the supply frequency, (100Hz in the UK), and falls to zero

between peaks. This means that the supply would exhibit 100% voltage ripple without

the inclusion of the smoothing capacitor. In addition to this the relatively long

duration between peaks means that the smoothing capacitor needs to be large to "hold

up" the supply and achieve an acceptably low amount of ripple.

We can estimate the size of smoothing capacitor required to obtain a particular

percentage of ripple. This is done by assuming that a constant current is drawn from

the capacitor over the 10ms time between charging pulses.

For example:

A 10kW 10kV supply must supply 1A average current. If we are prepared to accept

10% ripple, then the voltage across the smoothing capacitor is permitted to fall by

1kV over the 10ms duration between charging peaks. We can use the equation:

C = I x t / V

to find the required smoothing capacitance.

C = 1 x 0.01 / 1000 = 10 uF

This is a big capacitor which implies high cost. The 500 Joules of energy that it stores

are also highly dangerous. Clearly a trade-off exists between voltage ripple, and the

size and cost of the smoothing capacitor. Fortunately the demands made of the

smoothing capacitor and the resulting voltage ripple can both be reduced by choosing

a more elaborate supply arrangement.

3-pulse rectifier supply

The circuit below shows a simple 3-phase HVDC supply:

This design uses 3 independent step-up transformers followed by a 3-pulse rectifier

and a smoothing capacitor.

This 3-pulse rectifier essentially consists of 3 identical half-wave supplies feeding

into one smoothing capacitor.

Since the phases are spaced at 120 degrees relative to each other, then the capacitor

sees 3 charging pulses during each cycle of the mains supply.

The DC output voltage from this arrangement is also equal to 1.41 times the RMS

voltage rating of the transformer.

There are two advantages of this arrangement compared to the single phase supply

described previously. Firstly, the duration between charging pulses is now only

6.67ms instead of 10ms when used on a 50Hz supply. This means that the smoothing

capacitor does not need to be as big because it does not need to hold up the voltage for

so long. Secondly, the output voltage from this 3-pulse rectifier does not fall right

down to zero between pulses. This is because the 3-phases overlap slightly, and the

voltage ripple is actually 50% if no smoothing capacitor is used.

Clearly we are heading in the right direction by reducing the time between charging

pulses, and by reducing the "un-smoothed" ripple. Both of these things reduce the

demands on the smoothing capacitor. This reduces the system cost, and ultimately

will give superior performance.

Although the 3-pulse rectifier circuit is superior to a single phase supply, I would

not recommend actually building this supply for a number of reasons. The 3-pulse

rectifier only uses 3 HV diodes, so it is simple and cheap, but it is not very efficient

because it only makes use of the positive half-cycles from each transformer.

Secondly, the fact that it only uses the positive cycles from each transformer implies

an asymmetric loading on the secondaries of each transformer. This DC current

component is undesirable as it can result in saturation of the transformer cores.

A far more efficient supply can be built by using only 3 more diodes…

6-pulse rectifier supply

The circuit below shows a 3-phase supply using a 6-pulse rectifier:

This design uses 3 independent step-up transformers followed by a 6-pulse bridge

rectifier and a smoothing capacitor.

This 6-pulse rectifier is like a "full-wave" version of the 3-pulse design shown above.

Since both positive and negative half-cycles are used from all 3 phases, the capacitor

now sees 6 charging pulses during each cycle of the mains supply

The DC output voltage from this arrangement is 73% higher than that obtained from

the 3-pulse and single phase designs, because the 6-pulse rectifier extracts the

maximum phase-to-phase voltage. A 73% increase in voltage implies a tripling of

the energy in the Tesla Coil's primary capacitor, just for the cost of 3 additional HV

diodes ! The output voltage for this arrangement is equal to 2.45 times the RMS

voltage rating of the transformer.

There are a number of other advantages of this 6-pulse arrangement compared to the

two supplies discussed previously. Firstly, the duration between charging pulses is

now only 3.33ms with a 50Hz supply. This means that the size of the smoothing

capacitor can be reduced again, because it does not need to hold up the voltage for so

long. Secondly, the output voltage from this 6-pulse rectifier only falls to 86%

between peaks. This is because the 6 pulses overlap considerably, and the ripple is

only 14% without any smoothing capacitor. The reduced ripple means that the 6-pulse

supply could be used for Tesla Coil purposes without requiring any smoothing

capacitor.

Eliminating the smoothing capacitor represents a significant cost reduction in the

HVDC supply, and also removes a potentially dangerous source of stored energy from

the system. For this reason the author recommends the 6-pulse HVDC supply for DC

Tesla Coil use.

12-pulse rectifier supply

The process of using more charging pulses per supply cycle can be taken further in

order to reduce the "un-smoothed" ripple at the output of the supply. The circuit

below shows a more elaborate supply using a 3-phase supply and a 12-pulse rectifier:

This design uses 6 separate step-up transformers followed by a 12-pulse rectifier and

a smoothing capacitor.

The top half of the circuit is the same as the 6-pulse rectifier described above. (The

secondary windings of the three transformers are connected in Star Y configuration.)

The bottom half of the circuit is basically another 6-pulse rectifier, however the

secondary windings of these transformers are connected in Delta configuration.

This has the effect of shifting the phase of the bottom rectifier pulses by 30 degrees so

that they interleave perfectly between the pulses from the top rectifier.

When the outputs from the two 6-pulse rectifiers are combined, the smoothing

capacitor sees a total of 12 charging pulses during each cycle of the mains supply !

The DC output voltage from this arrangement is also equal to 2.45 times the RMS

voltage rating of the transformer, so there is no voltage gain in moving from the 6-

pulse arrangement to the more complex 12-pulse arrangement. However the duration

between charging pulses is now only 1.67 ms with a 50Hz supply, making life very

easy for any smoothing capacitor, if one is required at all.

The output from the 12-pulse rectifier only falls to 97% of its maximum voltage

between peaks. This equates to a ripple of only 3.5% compared to 14% for the 6-pulse

design above. Such a low ripple percentage makes this arrangement more than

adequate for our application without employing any smoothing capacitor at all

It should be realised that the lower 3 transformers in the circuit above, need to have

their secondary windings rated at 1.73 times the voltage of the upper 3 transformers.

This is because the secondary windings of the lower transformers are connected in

Delta configuration, and the upper ones are connected in Star (Y) configuration. If the

output voltages of the lower transformers are not scaled up accordingly, the 12-pulse

rectifier circuit will not function correctly.

Although the 12-pulse rectifier represents a technically elegant HVDC supply, with

minimal ripple, and no smoothing capacitor, the Tesla Coil designer must consider

whether the added cost, complexity (and weight) can be justified by the low ripple at

the DC output. For most applications the moderate ripple from the 6-pulse

arrangement is likely acceptable, but if you happen to come across a surplus 12-pulse

HVDC supply for the right price...

(Note that dedicated 3-phase transformers could be used instead of three discrete

transformers in most of the circuits above. However careful attention must be paid

to the way in which the 3 HV secondary windings are connected together. Most of

the circuits show here require that they are connected in star configuration as

access is required to the neutral for them to work correctly. Whether using separate

transformers or a single 3-phase unit, always pay attention to the dots next to the

windings in these circuits to ensure correct phasing.)

Scattered Radiation

Is radiation which arises from interactions of the primary radiation beam with the

atoms in the object being imaged. Because the scattered radiation deviates from the

straight line path between the X-ray focus and the image receptor, scattered radiation

is a major source of image degradation in both X-ray and nuclear medicine imaging

techniques. When X-ray radiation passes through a patient, three types of interactions

can occur, including coherent scattering (coherent scatter), photoelectric absorption

and Compton scattering . Of these three events, the great majority of scattered X-rays

in diagnostic X-ray imaging arise from Compton scattering.

In coherent scattering, the energy of the primary X-ray photon is first completely

absorbed and then re-emitted by the electrons of a single atom. Because no net energy

is absorbed by the atom, the re-emitted X-ray has the same energy as the original X-

ray, however the direction of re-emission is totally arbitrary.

In photoelectric absorption, the energy of the X-ray photon is completely absorbed as

it ejects a tightly bound electron from one of the atom's inner shells. The excess

energy of the photon over that of the binding energy of the electron is carried off as

kinetic energy by the ejected electron. Low energy characteristic radiation is

generated as an electron from an outer shell falls into the vacated lower shell.

Finally, in Compton scattering, the interaction can be considered as a collision

between a high energy X-ray photon and one of the outer shell electrons of an atom.

This outer shell electron is bound with very little energy to the atom and essentially

all of the energy lost by the X-ray photon in the collision is transferred as kinetic

energy to the electron, and the electron is ejected from the atom. Because energy and

momentum are both conserved in this collision, the energy and direction of the

scattered X-ray photon depend on the energy transferred to the electron. When the

initial X-ray energy is high, the relative amount of energy lost is small, and the

scattering angle is small relative to the initial direction. When the initial X-ray energy

is small, the scattering is more isotropic in all directions. At X-ray energies on the

order of 1 MeV (the energy range used in radiation therapy), the scattering is mostly

in the forward direction. At X-ray energies of 100 keV (the diagnostic imaging range),

the scattering is more isotropic.

The relative probability of the three types of interactions for different materials is

shown in Fig. 2. It is seen that in the diagnostic imaging range, near 100 keV,

Compton scattering comprises the great majority of interactions for normal tissues in

the body. The probability of photoelectric interactions increases as the substance

atomic number increases (going from water to bone for example) and as the X-ray

energy decreases. Coherent scattering is seen to be a very small fraction of the total

number of scattering events.

Scattered X-rays that arise from Compton scattering constitute a serious problem in

diagnostic imaging. Although the scattered X-ray photons are nearly isotropic in

direction at diagnostic energies, the scattered X-ray detected in the image are

primarily forward directed and thus have energies and angles of incidence near those

of the primary X-rays. Thus, these scattered X-rays cannot be completely removed by

the use of antiscatter grids or energy filters. The residual scatter reduces radiographic

contrast in X-ray imaging and contributes to image intensity distortion in computed

tomography CT .

Spectral sensitivity Film / Screen

Introduction

A film used in an X-ray cassette must have a spectral sensitivity that is matched to the

emission spectrum of the intensifying screen. Light emitted from an intensifying

screen in general can be either of two types; a continuous spectrum, as in the case of

CaWO0, or a band spectrum, as in the case of Gd2O2S:Tb (Lanex screens) (Fig.1).

A standard silver halide film will be sensitive to light up to a wavelength of 520 nm,

but will be almost insensitive to most of the light emitted by a gadolinium oxisulfide

screen (540 nm).

For this type of screen another type of film must be used. This type is called

orthochromatic film and is made sensitive to the green light from the screen by a

sensitizing dying agent in the emulsion that absorbs the green light and then transfers

the energy to the silver halide grains.

The spectral output of the phosphor must be matched to the response of the film (Fig.

2). Calcium tungstate screens emit blue light of continuous spectrum with a peak

wavelength at about 430 nm. The term "blue screen" refers both to the screen itself

and to the blue sensitive film used together with the CaWO4 screen. Rare earth

screens emit light in narrow lines with strong peak(s) in the green part of the spectrum

but smaller ones also in the blue, blue-green and yellow regions. The term "green

screen" may be used. It is absolutely necessary to use green sensitive film with these

screens to make sure that useful transmitted radiation is not lost.

Types of film and their sensitivity spectra

Monochromatic 300 - 500 nm

Orthochromatic 300 - 520 nm

Orthochromatic 300 - 580 nm

Orthochromatic Long 300 - 615 nm

Panchromatic sensitive to all wavelengths 300 - 760 nm

Thermo luminescent Dosimeter TLD

The thermoluminescent dosemeter (TLD) is designed to measure doses from X-, beta

and gamma radiations in terms of the radiation quantities specified by the Health and

Safety Executive (HSE). The TL dosimetry service is approved by the HSE under

Regulation 35 of the Ionising Radiations Regulations 1999.

The dosemeter consists of two thermoluminescent detectors containing the radiation-

sensitive material lithium fluoride. The detectors are located in a plate which is

identified uniquely by means of an array of holes. The lithium fluoride stores the

energy it receives from ionising radiations until it is heated during processing (in this

case to about 250°C) when the energy is released as light. The amount of light

released is proportional to the radiation dose. The plate is supplied to the wearer in a

plastic wrapper which protects the detectors from light and contaminants. This bears

the wearer's name, an establishment code, the expiry date and an optional personal

identifier for each employee, e.g. department name or a works number. If the name is

not required a serial number is printed instead. The wearer places the wrapped plate in

a plastic holder,which is supplied by the NRPB on permanent loan and is available

with safety pin or clip attachments. We also provide 'wear and care' cards for each

member of staff. These are designed to help users understand how and why they

should wear the dosemeter.

The dosemeter measures two quantities. The first is the personal dose equivalent Hp

(10), which is often referred to as the 'whole body' dose which results from

penetrating radiation. It is measured by the detector behind the domed part of the

holder. The second quantity is the personal dose equivalent Hp (0.07) which is an

assessment of the dose equivalent to the skin from both weakly and strongly

penetrating radiations. This is measured by the detector behind the circular window.

TLD technical specification

Detection x rays and gamma rays beta particles

Dose range

measured 0.05 mSv to 10 Sv 0.05 mSv to 10 Sv

Energy range

detected

10 keV to 10 MeV for Hp (0.07)

15 keV to 10 MeV for Hp (10)

700 keV to 3.5 MeV (Emax) for

Hp (0.07)

Periods of use 2, 4, 8, 13 weekly (calendar

issue periods are also available)

2, 4, 8, 13 weekly (calendar

issue periods are also available)

Special features of the TLD

Tissue equivalence

The detectors absorb radiation energy in the same way and to the same extent as

human tissue. This enables us to evaluate doses of complex mixtures of radiations in a

simple and straightforward manner, thus keeping errors of measurement to a

minimum.

Life span

The detector is capable of retaining the stored dose information for extended periods

before assessment. Even in conditions of relatively high temperature (40°C) and high

humidity (up to 100%), the information can be stored for up to one year. Issue periods

of up to 13 weeks can be offered thus keeping the cost of monitoring low.

Reassessment of TLD

TL glow curves of all dosemeter readings are kept for at least five years. This allows

retrospective investigation in the event of a customer query. The glow curves for

dosemeters with assessments in excess of 15 mSv are all checked. For doses over 25

mSv, it is possible to verify the original assessment using a special technique, at no

extra cost.

Image 1& 2 Typical TLD holder & card

Image 3 TLD Badges

Image 4 TLD Ring types

The TLD

Card

The TLD

Holder

Image 4 A selection of TLDs from Canada

Tomography

Tomography otherwise known as body section radiography, planigraphy,

laminography or stratigraphy, is the process of using motion of the X-ray focal spot

and image receptor (e.g. film) in generating radiographic images where object detail

from only one plane or region remains in sharp focus

Fig. 1. a and b

Two lateral tomograms of the temporal bone 6 mm apart, acquired with hypocycloidal

tomography. The more external tomogram (a) shows the midpart of the

temporomandibular joint, and the more medial tomogram (b) shows an exostosis in

the anterior part of the external auditory canal (arrow).

Fig 2

Tomography, Fig. 2

The motion of X-ray tube and film in linear tomography.

Details from other planes in the object which would otherwise contribute confounding

detail to the image, are blurred and effectively removed from visual consideration in

the image. A variety of tomography techniques have been developed, which differ

primarily in the manner in which the X-ray source and film move.

Linear tomography is one of the most basic techniques (Fig. 2). As the tube and film

move from the first position to the second, all points in the focal plane project to the

same position on X-ray film. Thus, points a, b and c project to points a', b' and c' in

the first position and a", b" and c" in the second position. Points above or below the

focal plane do not project to the same film positions and are blurred. By changing the

relative motion of the film and tube, the focal plane can be adjusted upward or

downward.

In addition to linear tomography, other types of tube and film motion have been used.

These motions include circular, elliptical, figure-8, hypocycloidal, trispiral (Fig. 3).

Each of these motions has advantages regarding the way in which out of plane

structures are blurred. For example, a linear structure which is aligned with the linear

motion of a linear tomograph, will not appear blurred, except at the ends, whereas

such a structure will be blurred by the circular motion of a circular tomograph.

Fig 3

Tomography, Fig. 3

Alternative tomographic motions. circular tomography and hypocycloidal

tomography.

Circular tomography, a tomographic method where the X-ray focus and the film cassette are moved in

circular patterns. The X ray tube and cassette holder are mechanically connected and

move in a pattern as demonstrated in Fig.1. As can be seen from the figure, the film

cassette does not rotate along its path.When grids are used, the grid lines must follow

the rotation in order to prevent grid cut off. The advantage of circular tomography is

that a uniform body section thickness is obtained in the image. The disadvantages are

the long exposure time and the complex design of the equipment.

Hypocycloidal tomography

Hypocycloidal tomography,

tomography in which the X-ray tube and film move in a hypocycloidal path.

Conventional tomography can be made using several movement patterns for the X-ray

tube and the film. The common linear movement is mechanically easy to produce but

will give rise to rather thick tomographic sections and a short blurring path (the length

of the tomographic section). If thinner sections and longer blurring paths are required,

more complex movements are needed. Circular motion will for the same angulation of

tube and film produce three times longer blurring paths than the linear motion and

thinner sections. However, artefacts can be generated for circular-shaped objects in

the tomographic plane. Spiral movement, which is a combination of circular and

radial movement, will overcome the artefacts, but requires that the tube (and film)

speed decreases when the tube is spiralling out from the centre of the spiral. This is

difficult to achieve mechanically.

The hypocycloidal movement is also a combination of a circular and radial movement

(Fig. 1). The pattern can be produced by letting an inner gearwheel rotate inside

another gearwheel with teeth on the inside. The proportion of "teeth" inner/outer

wheel is 2/3. If the tube and film support is connected to the centre of the inner wheel,

it will describe a hypocycloidal movement. This movement is fairly easy to achieve

mechanically and performs superior to all others. It will produce thin sections with a

blurring path five times longer than for linear movement with the same angulation. No

object should present a hypocycloidal shape, so virtually no artefacts will be produced.

The only disadvantage is that the tomographic section produced is extremely thin,

which imposes the need for very high precision with regard to the film position.

Notes - Zonography - a form of tomography where the tomographic angle is small,

on the order of 10, resulting in a thick plane of focus. The technique is sometimes

used to better delineate suspected pathology.

Pantomagraphy - a special tomography technique where panoramic roentgenograms

of curved surfaces are obtained by rotating the X ray tube and film-screen holder

around the patient, who is usually in a sitting position (Fig.1). The film holder, which

is much longer than the film, has a protective lead front with a narrow slit. The film is

exposed through this slit starting from one end. The film moves across the slit as the

X-ray tube and film holder rotate around the patient. The result is a PA image of a

curved surface, e.g. the mandible, flattened out on the two-dimensional film. In

dentistry radiology, the technique is also called orthopantomography and is there still

in much use, while other conventional tomographic techniques have been mostly

replaced by computed tomography CT .

Transformers

A transformer consists of two coils (often called 'windings') linked by an iron core, as

shown in figure 1. There is no electrical connection between the coils, instead they are

linked by a magnetic field created in the core.

Transformers are used to convert electricity from one voltage to another with minimal

loss of power. They only work with AC (alternating current) because they require a

changing magnetic field to be created in their core. Transformers can increase voltage

(step-up) as well as reduce voltage (step-down).

Alternating current flowing in the primary (input) coil creates a continually changing

magnetic field in the iron core. This field also passes through the secondary (output)

coil and the changing strength of the magnetic field induces an alternating voltage in

the secondary coil. If the secondary coil is connected to a load the induced voltage

will make an induced current flow. The correct term for the induced voltage is

'induced electromotive force' which is usually abbreviated to induced e.m.f.

The iron core is laminated to prevent 'eddy currents' flowing in the core. These are

currents produced by the alternating magnetic field inducing a small voltage in the

core, just like that induced in the secondary coil. Eddy currents waste power by

needlessly heating up the core but they are reduced to a negligible amount by

laminating the iron because this increases the electrical resistance of the core without

affecting its magnetic properties.

Transformers have two great advantages over other methods of changing voltage:

1. They provide total electrical isolation between the input and output, so they

can be safely used to reduce the high voltage of the mains supply.

2. Almost no power is wasted in a transformer. They have a high efficiency

(power out / power in) of 95% or more.

Types of Transformer

Mains Transformers

Mains transformers are the most common type. They are designed to reduce the AC

mains supply voltage (230-240V in the UK or 115-120V in some countries) to a safer

low voltage. The standard mains supply voltages are officially 115V and 230V, but

120V and 240V are the values usually quoted and the difference is of no significance

in most cases.

To allow for the two supply voltages mains transformers usually have two separate

primary coils (windings) labelled 0-120V and 0-120V. The two coils are connected in

series for 240V (figure 2a) and in parallel for 120V (figure 2b). They must be wired

the correct way round as shown in the diagrams because the coils must be connected

in the correct sense (direction)

:

Most mains transformers have two separate secondary coils (e.g. labelled 0-9V, 0-9V)

which may be used separately to give two independent supplies, or connected in series

to create a centre-tapped coil (see below) or one coil with double the voltage.

Some mains transformers have a centre-tap halfway through the secondary coil and

they are labelled 9-0-9V for example. They can be used to produce full-wave rectified

DC with just two diodes, unlike a standard secondary coil which requires four diodes

to produce full-wave rectified DC.

A mains transformer is specified by:

1. Its secondary (output) voltages Vs.

2. Its maximum power, Pmax, which the transformer can pass, quoted in VA

(volt-amp). This determines the maximum output (secondary) current, Imax...

3. ...where Vs is the secondary voltage. If there are two secondary coils the

maximum power should be halved to give the maximum for each coil. Its

construction - it may be PCB-mounting, chassis mounting (with solder tag

connections) or toroidal (a high quality design).

Turns Ratio and Voltage

The ratio of the number of turns on the primary and secondary coils determines the

ratio of the voltages...

where Vp is the primary (input) voltage, Vs is the secondary (output) voltage, Np is the

number of turns on the primary coil, and Ns is the number of turns on the secondary

coil.

Power and Current

The very small power loss in a transformer means that we can assume that power in =

power out. Power = voltage x current, so we can use this to show that the current ratio

is the inverse of the voltage ratio. For equal power the current increases as the voltage

decreases

so...

where Ip is the primary (input) current, and Is is the secondary (output) current.

The current in the primary coil Ip is determined almost entirely by the current Is drawn

by the load connected to the secondary coil. With no load connected, Is = 0 and Ip is

very small indeed because the alternating magnetic field created by the primary

current induces a voltage in the primary coil which almost exactly matches the supply

voltage. If a load is connected current will flow in the secondary coil, creating a

magnetic field which opposes and partly cancels out the field created by the primary

coil. The resulting weakened field induces a smaller voltage in the primary coil to

oppose the supply voltage and this means that a larger primary current flows.

Unsharpness

There are three pricipal types of unsharpness associated with traditional imaging

methods

Motion

Geometric

Photgraphic

Note CR and DR imaging have other causes of unsharpness

Unsharpness,

a quantitative measure of the loss of edge detail which is due to geometric properties

of the object and imaging system and not due to image noise or X-ray scatter. It is

usually expressed as the width of the band of changing density or brightness arising

from a sudden change in the intensity of the radiation incident on the film or

fluorescent screen. From this definition it can be understood that unsharpness and

resolution are different concepts. It is possible for an edge to be "spread" by one of

many factors, and at the same time for two such edges to be resolved in the image.

The factors which contribute to the total image unsharpness include geometric

unsharpness, movement unsharpness, absorption unsharpness, image receptor

unsharpness, and parallax unsharpness. The various unsharpness factors all contribute

to the observed unsharpness of structures in an image. However, the quantitative

manner in which the factors combine is in general complicated and is not completely

understood. It is known from observation that the total unsharpness is not the direct

sum of the contributing factors. In general, it appears that the total image unsharpness

is dominated by the unsharpness of the largest individual factor.

Motion artefact,

artefact occurring whenever image acquisition takes longer than the time over which

physiological motion occurs in the body region of interest. Motion artefacts are

usually not a problem in imaging the brain and the extremities, except when the

patient cannot lie still during the examination, but they can be prominent when

imaging the trunk. Typical periods over which physiological motion occurs and an

approximate severity scale for the motion effects on image quality are given in table 1.

Motion artefact, Table 1

Examples of physiological motion, its duration and effect on imaging

Body region Type of motion Severity of effect Period of motion

Brain Cerebrospinal fluid (cardio-sync.) + 100 ms

Blood flow + 100 ms

Spine Cerebrospinal fluid (cardio-sync.) + - ++ 100 ms

Neck Glutition + suppressible

Respiration ++ 4 s

Blood flow ++ 100 ms

Thorax Respiration +++ 4 sec

Cardiac motion ++++ 50 ms (sys.) - 400 ms (diast.)

Blood flow +++

Upper abdomen Peristalsis ++ 10 s

Respiration ++ - +++ 4 s

Blood flow ++ 100 ms

Lower abdomen and pelvis Peristalsis + 30 s

Blood flow + 100 ms

Extremities Blood flow + 100 ms

The best way of suppressing image motion artefacts is to acquire data faster than the

typical periods given in Table 1. However this does not work for all imaging

modalities. For different imaging techniques, and particularly for MRI, various

ingenious ways have been devised to suppress motion artefacts such as cardiac gating,

respiratory gating and motion compensation.

Photgraphic unsharpness

Photographic unsharpness factors in Film screen radiography

Film emulsion grain size

thickness of the emulsion layer

single vs double emulsion film

cross-over in double emulsion

Screens thickness of the phosphor layer

size of the phosphor crystals

reflective layer

absorbing layer

dye tint

Screen unsharpness, the contribution to image blurring or unsharpness due to

spreading or diffusion of light within the intensifying screen and between the screen

and film surfaces. Because the screen has a finite thickness, the X-ray absorption

event which generates the emission of light within the screen may occur at some small

distance from the film. The light diverges from that point and has spread a small

distance, related to the screen thickness, by the time it reaches the film surface

Parallax unsharpness, an image unsharpness seen only in double emulsion film. In

principle, there is an image in both emulsions, separated by the thickness of the film

base, about 0.1-0.2 mm. If the film is looked at from an angle, these two images do

not overlap exactly causing parallax unsharpness. Its influence to total image

unsharpness is negligible.

Parallax,

the apparent displacement of an object when viewed from two different angles, e.g.

when observing an object first with the right eye and then with the left eye (Fig.1). In

Figure 1, the apparent position of object A with respect to object B changes when the

view shifts from one eye to the other. Due to the shorter distance to object A than to

object B, the convergent angle from object A (a) is larger than that from object B (b).

The difference in angles (a - b) is called the angle of instantaneous parallax

Geometric unsharpness,

unsharpness in the image caused by the fact that the X-rays are emitted from an area

rather than from a point. Regions at the edges of an object will be formed in which the

X-ray intensity will be gradually increasing (or decreasing), causing unsharpness (see

geometric magnification (I), Fig. 1). These regions are generally referred to as

penumbra. The magnitude of the penumbra is dependent on the focal spot size and the

ratio focus-object distance/focus-film distance.

Geometric magnification,

the (theoretical) magnification in an X-ray image that occurs when the focal spot is

assumed to be a point and not an area. The magnification of an object is easy to

calculate, given the focus-object and focus-film distances, respectively, and assuming

that the focal spot is a point (Fig. 1, left). The magnification M, is then:

M = d/c = (a+b)/a

However, if the actual size of the focal spot is taken into account, the geometry is not

the same. The image will now be slightly more magnified having, however, a more

diffuse edge due to the penumbra present . The magnification is now:

M = [(a+b)/a] + {[(a+b)/a] - 1}(f/c)

where f is the diameter of the focal area. When the focal spot size is accounted for, the

magnification is referred to as "true magnification".

Turbidity

image unsharpness due to radiation scattered by the photographic emulsion

X-Ray Spectrum

The distribution of the number of X-ray photons with a certain energy as a function of

the energy is called the photon spectrum or quantum spectrum of the radiation.

Graphs of spectra usually show only relative or normalized numbers of photons, as

this is the only information needed for most applications, e.g. in order to calculate

contrast values in X-ray imaging. By multiplying the photon numbers with the

associated values of the individual photon energy, another type of spectrum can be

obtained: the intensity spectrum. This spectrum directly shows, which fraction of the

total beam intensity stems from the various photon energy ranges of the spectrum, and,

therefore, allows an immediate calculation of total signal intensity (or energy) by

mathematical integration. The graphical representation of the integral is simply the

area under the curve. Both types of spectra look very similar. For qualitative

discussions of spectral effects, it usually is rather unimportant which type of spectrum

is used, but for a quantitative analysis, it is essential.