Aside from ultrasound, a digital radiography suite is going to be the most expensive upgrade for the private practitioner
in the realm of diagnostic imaging. However, this expense will be well worth the investment from a diagnostic imaging standpoint.
After more than a century of film and film screen imaging, the backdrop of diagnostic radiology has changed from hanging films
on view boxes to LCD monitors.
X-ray Production Overview
X-rays are produced by the acceleration of electrons from a cathode source (made of tungsten) to an anode target (made of 90%tungsten/10% rhenium and can be a stationary or rotating target) with both collisional (>69.5 kVp) and
decelerating interactions resulting in characteristic and to a greater extent, Bremsstrahlung radiation respectively. The purpose of a rotating anode is to spread the heat produced during an exposure over a large area
of the anode. The cathode and anode for standard x-ray tubes are made out of Tungsten, which has a melting point of 3400° C, and a Z of 74 and an inner k-shell electron binding energy of 69.5 keV. The cathode
(-) is a coiled filament that measures 1.5 mm in diameter and 10 to 15 mm in length. The electron cloud is created when a
low current is applied to the cathode during tube "warm-up" prior to exposure. This electron cloud is held together by a
strong negative charge, surround the cathode filament, heated within the focusing cup. The focusing cup is electrically connected to the filament circuit and provides a voltage of up to 10 V to the filament,
which produces a current up to about 7 Amps. The mA determines the number of electrons boiled away from the tungsten cathode
as the filament is heated secondary to electrical resistance within the filament circuit by a process called thermionic emission. As one increases the mA, the temperature of the cathode increases so that more electrons are released through the process
of thermionic emission.
The size of the cathode is one factor that will determine the degree of penumbra (edge unsharpness) present on the final radiographic image. Typically x-ray tubes used for magnification will have small
focal spots on the order of 0.3 × 1.2 mm vs. the larger focal spots typical for the small animal veterinary products currently
on the market, (1.0 × 2.0 mm).
In addition to actual focal spot size, the effective focal spot size will be smaller (length only and not width) due to the
angle of the anode due to the line-focus principle. The projected or effective focal spot may differ by 50% from the actual nominal size. The higher the mA station, the larger
the electron bloom results, thereby increasing the actual size of the electron cloud. When using small focal spots, it is
best to try to minimize the mA (usually mA limited on generator controls). The effective focal spot is the size of the focal
spot as projected down the central axis of the primary x-ray field. The relationship between the actual and effective focal
spot lengths is as follows:
Effective focal length = Actual focal length x sin(β ), where β = anode angle.
The major trade-offs to consider with regards to anode angle include: smaller anode angle (7 to 9 degrees) = small effective
focal spot (better spatial resolution); limit to size of field coverage, and poor power loading ability of the anode due to
the higher heat load placed focally on the anode. This type of anode angle would be used for limited field coverage specialty
imaging such as neurovascular imaging, etc. Conversely, a larger anode angle (12 to 15 degrees) would equate with good field
coverage, larger effective focal spot (increased penumbra) and good power loading (higher techniques or heat unit capability
due to spread of the actual focal spot). The power loading or intensity of heat units is dependent upon the cathode filament
length so that a longer filament length will decrease the intensity of the heat units placed per unit area on the anode (good
power loading), whereas a small cathode filament length results in higher heat units per unit area on the anode (poor power
loading). Typically tubes are described based on the effective focal spot length and width with common lengths being 0.3,
0.6, 1.0, 1.2 and 2.0 mm.
There are anode-cathode differences in intensity noted particularly in large field of view images with low kVp technique.
These intensity differences are called the heel effect and is based on the fact the there is a reduction in the intensity of the primary x-ray beam along the anode side of the
field of view. This is because the x-rays that have to traverse the anode side of the angled anode are more readily attenuated
by anode material. Because of this the heel effect is less noticeable for higher kVp techniques, longer tube-film distances,
smaller fields of view and small anode angles (7 to 10 degrees).
X-ray generators that are commonly used in veterinary medicine include: single phase (1φ), half-wave and full-wave rectified
(1φ2 ), three phase (3φ6 or 3φ12 )and high frequency generators. In single-phase half-wave generators, the incoming 60 Hz line is divided into positive and
negative cycles for the incoming 120 Volts. A single-phase generator that is half wave rectified only takes advantage of
the positive phase of the incoming pulse, therefore the fastest time for this type of generator will be 1/60th of a second. For a fully rectified, single-phase generator there is still 100% ripple as seen with the previous generator,
but the rectifiers will allow the operator to take advantage of both the negative phase (by flipping it to a positive phase
cycle) and the positive phase of the incoming 60 Hz alternating current. This means that the fastest time one can use is
1/120th of a second. If there is patient motion on the radiograph, one can count the number of images present (since there is 100%
ripple) and determine the timer used for the radiographic exposure. If one knows the type of generator, one can account for
whether each pulse is 1/60th of a second or 1/120th of a second and determine the timer.
Example: If you have a radiograph and can count 6 images, calculate the timer used for a single-phase half wave rectified
and full wave rectified generator.
Half Wave, Single Phase Generator
1 pulse = 1/60th sec; 6 pulses (1/60th sec/1 pulse) = 6/60 = 1/10 of a second timer used.
Full Wave, Single Phase Generator
1 pulse = 1/120th sec; 6 pulses (1/120th sec/ 1 pulse) = 6/120 = 1/20 of a second timer used.
Average X-ray energies are increased when comparing single-phase generators versus three phase generators. In a three-phase
generator, the average kilo voltage is higher depending on the type of generator and can approach the actual kVp selected
for a twelve-impulse, three phase generator, because there is virtually 3 to 5% ripple. The x-ray intensity is therefore
higher and the average x-ray beam energy is also higher for three phase generators. For a 3φ6 or 3φ12 one will have to use one half the mAs as with a 1φ2 in order to obtain the same optical density on the radiographic film.
kVP (kilo voltage peak) will control the x-ray penetration, while the mA ultimately controls the number of electrons boiled
off of the focal spot cathode, thereby impacting the number of x-rays produced. This is important when considering contrast
within a radiographic image due to the basics in x-ray attenuation within the patient. For a high contrast image (defined
as stark transitions between dark gray and light gray or white) one would use a low kVp, high mA technique where as for a
latitude (longer gray scale image) one would use a high kVp, low mA technique.