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X‐Rays

Introduction

X‐rays routinely are used to non‐invasively examine internal anatomy. The technique is called Radiography. The image created using x‐rays is called a radiograph. Many people confuse the terms “x‐ray” and “radiograph.” X‐rays are a type of energy, whereas radiographs are images.

Radiography is a part of the medical specialty called Radiology. Radiology encompasses all of diagnostic imaging, including ultrasound, computed tomography (CT), magnetic resonance imaging (MRI), and scintigraphy. A major advantage of radiography over the other imaging modalities is that a large volume of the patient can be viewed with one image. A major disadvantage is that many structures are superimposed, making interpretation of the images sometimes confusing.

Our goal in radiography is to make images that display the greatest amount of usable information. This requires an understanding of x‐rays; what they are, how they are made, and how they can be used safely to “look inside” our patients.

X‐rays

X‐rays were discovered in 1895 by Wilhelm Conrad Röntgen. Röntgen was a German scientist who won the first Nobel Prize in physics for his discovery (Figure 1.1). He used the symbol “X” to name the new rays because they were a type of radiation that was unknown at that time. Radiation is energy that spreads out in all directions as it moves away from its source.

X‐rays are a type of electromagnetic radiation (EMR). There are many types of EMR, including radiowaves, microwaves, and visible light (Figure 1.2). EMR behaves both like particles and like waves, sometimes described as a stream of particles that travels in a wave‐like pattern. EMR particles are called photons. Photons are packets of energy with no mass, no charge, and move at the speed of light. The energy of a photon is directly related to wavelength; the shorter the wavelength, the greater the energy. Wavelength is directly related to frequency; the higher the frequency, the shorter the wavelength (and the higher the frequency, the greater the energy).

In the EMR spectrum, x‐rays are relatively short wavelength, high frequency, and high energy. They are able to penetrate many materials that are opaque to visible light (Box 1.1). X‐rays also are ionizing, which means they have sufficient energy to remove electrons from atoms and form ions. Ionization can destroy molecules and damage or kill living cells. Adherence to radiation safety guidelines is crucial when working with x‐rays (see the Radiation Safety section later in this chapter).

The reason we are able to use x‐rays to view internal anatomy is because x‐rays pass through some tissues more easily than others. Without differences in tissue transmission, radiography would not be possible. The ability of a tissue to block or attenuate x‐rays is called radiopacity or simply opacity (“radio‐” is assumed since we are discussing radiography). The greater the opacity, the fewer x‐rays will pass through.

(A) Photo of Wilhelm Röntgen, the German physicist who discovered x-rays. (B) Photo of the first radiograph made by Röntgen. — **Figure 1.1** Discovery of x‐rays. (A) Photo of Wilhelm Röntgen, the German physicist who discovered x‐rays.

Source: JdH/Wikimedia Commons/Public domain.

(B) Photo of the first radiograph made by Röntgen. It is an image of his wife’s hand. The ring on her finger is the first radiographic artifact (arrow).

Source: NASA.

Another key point in radiography is that a tissue or object will be visible in a radiograph only if it is adjacent to a material with a different opacity. There must be an opacity interface. An opacity interface is the boundary where two materials with different opacities meet. Without opacity interfaces there is no visible image in a radiograph. The greater the difference in opacities between materials, the stronger the opacity interface and the easier it will be to identify structures.

Schematic illustration of spectrum of electromagnetic radiation. — **Figure 1.2** Spectrum of electromagnetic radiation (EMR). Different types of EMR are displayed in this image. The type is determined by the energy of the photons. The shorter the wavelength, the higher the frequency, and the greater the energy. High energy EMR can be ionizing.

Source: trekandphoto/Adobe Stock.

As mentioned earlier, our goal is to produce radiographs that provide maximum usable information. Making quality radiographs requires adequate density, contrast, detail and minimal distortion. Each of these factors is discussed thoroughly later in this chapter. However, a brief introduction may be helpful here.

Density refers to the amount of blackness in the radiograph. The greater the radiographic density, the darker the image. Density is determined by the number of x‐rays that are recorded by the image receptor. The more x‐rays recorded, the darker that part of the radiograph. In most radiographs there are many levels of density. Each level is displayed as a shade of gray. Each shade of gray corresponds to an opacity. If there is too much density, the radiograph will be too dark. If there is too little density, the radiograph will be too light. In either case we won’t be able to differentiate structures in the image.

Contrast refers to the amount of difference between the levels of density in the radiograph. It is the degree of change between the dark areas and the light areas. High contrast means the radiograph is mostly black and white with few shades of gray in between. Some structures will not be visible because there are no shades of gray to display their opacities. Low contrast means there are many shades of gray between white and black. If there is too little contrast, structures with similar opacities will not be distinguished because their individual shades of gray will be too similar and they will visibly blend together.

Detail refers to the sharpness of change from one density to the next in the radiograph. The greater the detail, the sharper the edges of the structures in the image. Detail – or definition – determines how well we can see the borders of structures that are close together. If there isn’t enough detail, we won’t be able to distinguish the structures as separate.

Distortion occurs when the actual size or shape of an object is misrepresented in the radiograph. It results from poor positioning of the object in relation to the x‐ray beam and image receptor.

X‐ray production

For medical radiography, x‐rays are created inside a glass vacuum tube. This is accomplished by bombarding a metal target with high‐speed electrons. Inside the tube are a negatively charged cathode and a positively charged anode (Figure 1.3). The cathode contains a wire filament which can be heated with an electric current to “boil off” electrons (Figure 1.4). This process is called thermionic emission and it is similar to heating the filament in an incandescent light bulb. The number of electrons produced is controlled by the strength and duration of the current.

Schematic illustration of x-ray tube. — **Figure 1.3** X‐ray tube. This is a picture of a generic x‐ray tube showing the positively charged anode (+) and the negatively charged cathode (−) enclosed in a glass vacuum tube.

Schematic illustration of x-ray production. — **Figure 1.4** X‐ray production. This illustration depicts the filament in the cathode being heated to emit a cloud of electrons. The electrons (e⁻) are aimed at the anode by the focusing cup and accelerated by a high voltage. The electrons strike a target on the anode (black area) and their kinetic energy is converted into heat and x‐rays. The x‐rays produced vary in wavelength and energy.

The electrons collect in a “cloud” around the filament and are held in place by the focusing cup. The focusing cup concentrates the electrons into a well‐defined beam and aims them at the anode. When high voltage is applied across the x‐ray tube, a large negative‐to‐positive gradient is created that rapidly accelerates the electrons from the cathode to the anode. The electrons are traveling at over half the speed of light when they strike the anode.

The anode stops the electrons, converting their kinetic energy into heat and x‐rays (about 99% heat and 1% x‐rays). Heat is an unwanted by‐product of x‐ray production and must be dissipated. Methods of heat dissipation will be discussed shortly.

Electron bombardment of the anode produces x‐rays with energies that vary from near zero to the maximum energy of the electrons. The variations in energy are due to the way the electrons interact with the anode atoms. There are two types of interactions: characteristic and bremsstrahlung.

Characteristic x‐rays are formed when a high‐speed electron from the cathode collides with an atom in the anode target (Figure 1.5). An inner level electron is ejected from the atom, which produces an ion. Ions are unstable, so an outer level electron drops down to fill the void in the inner level. Because the binding energy of an outer electron is greater than that of an inner electron, the outer electron releases some energy in the form of an x‐ray. The energy of the x‐ray is equal to the difference between the outer level binding energy and the inner level binding energy and therefore is characteristic of the type of atom that was ionized. About 20% of medical x‐ray production is due to characteristic interactions (Figure 1.7).

Schematic illustration of characteristic x-ray production. — **Figure 1.5** Characteristic x‐ray production. 1. A high‐speed electron from the cathode (green e⁻) collides with an inner shell electron in an anode atom. 2. The anode electron is ejected, and the cathode electron continues in a new direction with less energy. 3. An outer shell electron gives up energy as an x‐ray and moves to fill the vacancy in the inner shell (N⁺= nucleus, e⁻= electron).

Bremsstrahlung x‐rays are produced when a high‐speed electron from the cathode passes near the nucleus in an anode atom. The positive charge from the nucleus pulls on the electron, slowing it down and altering its direction (Figure 1.6). “Bremsstrahlung” is a German word that means “braking radiation.” As it slows down, the electron loses energy which is emitted as an x‐ray. The energy of the x‐ray depends on how much the electron was slowed down. X‐ray energies vary from near zero (the electron barely slowed down) to the total energy of the electron (the electron was completely stopped). About 80% of medical x‐ray production is due to Bremsstrahlung interactions (Figure 1.7).

Schematic illustration of bremsstrahlung x-ray production. — **Figure 1.6** Bremsstrahlung x‐ray production. A high‐speed electron from the cathode (green e⁻) nears the positively charged nucleus (N⁺) in an anode atom. The electron slows down and changes direction. As it slows, the electron loses energy which is emitted as an x‐ray. The energy of the x‐ray is determined by how much the electron was slowed down.

Schematic illustration of spectrum of x-ray energies. — **Figure 1.7** Spectrum of x‐ray energies. Most of the x‐rays used for diagnostic imaging will range in energy from near zero to the maximum energy of the electrons used to produce them. This continuum of energies is due to Bremsstrahlung interactions and is shown in blue in the graph above. A small amount of medical x‐rays are characteristic of the material used to produce them, depicted by the green spikes on the graph.

X‐ray machine

Basic components of an x‐ray machine include the x‐ray tube assembly, the collimator, the electrical circuits, the operating console, the table, and the image receptor (Figure 1.8).

Schematic illustration of typical xEnd-ray machine. — **Figure 1.8** Typical x‐ray machine. The x‐ray tube assembly is mounted above the table, preferably on a slide for horizontal movement and on an adjustable column for vertical movement. The collimator is mounted to the x‐ray tube. The image receptor may be positioned atop or underneath the table top. In this picture, the image receptor is secured in a tray under the table, which is aligned with the collimator. The operating console is pictured on the right.

X‐ray tube assembly

The x‐ray tube assembly comprises the x‐ray tube and its metal housing (Figure 1.9). The metal housing protects and supports the tube and prevents unwanted x‐rays from escaping. It also encases the oil that surrounds the x‐ray tube to help dissipate heat and provide electrical insulation. Most x‐ray assemblies are mounted so they can be moved horizontally, vertically, and rotated to facilitate making radiographs.

The purpose of an x‐ray tube is to convert electricity into x‐rays. As described earlier, the cathode is the negative terminal in the x‐ray tube. It contains the filament and focusing cup. The filament is a coil of thin wire, usually made of tungsten. Tungsten is a very malleable metal with a high melting point (over 3400 °C) and a low rate of evaporation, making it ideal for thermionic emission. The focusing cup generally is made of molybdenum. The point on the anode that is struck by the beam of electrons is called the focal spot. Most x‐ray machines include two focal spots, a large one and a small one, with a separate filament for each.

The anode is the positive terminal in the x‐ray tube. This is where x‐rays are produced. A typical anode consists of a round, flat, metal disc made of molybdenum with a thin rim of tungsten. The high atomic number of tungsten (74) increases the likelihood of electron interactions. The anode rim is the target for the electron beam. It is beveled to create a slope, which is valuable for x‐ray beam geometry, as we will discuss later.

Heat is a major limiting factor in the production of x‐rays. To produce an adequate number of electrons, the tungsten filament must be heated to over 2200 °C. To accelerate the electrons, kilovolts of power must be applied across the x‐ray tube. To generate x‐rays, the target is heated to over 2500 °C during a single exposure. Excessive heating is a primary cause of x‐ray tube failure, and the x‐ray tube is one of the most expensive parts of an x‐ray machine. The most vulnerable part of the tube to heat overload is the focal spot. The larger the focal spot, the greater the heat tolerance. A large focal spot is desirable because more x‐rays can be produced; however, the larger the focal spot, the less definition in the radiographs.

Heat is conducted away from the focal spot by the metal in the anode disc, the oil surrounding the x‐ray tube, and the protective metal housing. To further dissipate heat, the anode is mounted on a high‐speed rotor which spins up to 10,000 rpm during x‐ray production. Spinning the anode effectively increases the area of the target that is bombarded by the electron beam and lessens the amount of heating in any one spot.

To avoid overheating, tube rating charts are provided by manufacturers. These charts define the maximum heat units a particular x‐ray tube can tolerate over time before it fails. Heat units (HU) are calculated based on the amount of power that is sent to the x‐ray tube (HU = kV × mA). You should never exceed 80% of the maximum heat units with any exposure. Anode cooling charts display the minimum time required for the anode to cool down before making another exposure.

X‐ray machines that have been idle for more than 6 hours are considered “cold.” A cold x‐ray tube should be warmed up prior to making a high exposure to avoid damaging the anode. To warm up a cold tube, close the collimator and set a low exposure technique (e.g., 60 kVp, 100 mA, 0.05 seconds). Make two exposures about 30–60 seconds apart. The tube is now warmed up and ready for use (Box 1.2).

Collimator

A collimator is an arrangement of x‐ray absorbers used to limit the size and shape of the x‐ray beam to a specific field‐of‐view (FOV). The FOV is the area to be irradiated. It is the part of the patient that is exposed to the primary x‐ray beam.

Collimators are mounted to the x‐ray tube assembly in the path of the x‐rays (Figure 1.10). Modern collimators contain adjustable shutters made of lead which can be manipulated to create the desired FOV. The smaller the FOV, the better the definition in the radiograph.

Schematic illustration of collimator. — **Figure 1.10** Collimator. In this illustration, the collimator is mounted to the x‐ray tube assembly in the path of the x‐rays. The side of the collimator is cut away to show the adjustable lead shutters inside. The shutters can be moved to change the size and shape of the FOV. The side of the protective housing is cut away to show the x‐ray tube. An added filter is visible between the collimator and x‐ray tube assembly, in the path of the x‐rays.

Inside many collimators are a light bulb and an array of mirrors which are used to project a visible light FOV that corresponds to the x‐ray beam FOV. This allows the radiographer to accurately configure the area to be irradiated.

Filtration

The purpose of filtration is to absorb low‐energy x‐rays and remove them from the x‐ray beam. Low energy x‐rays are unable to pass through the patient and therefore provide no useful diagnostic information. They do, however, increase the radiation dose to the patient. All x‐ray beams contain low‐energy x‐rays due to Bremsstrahlung interactions. The National Council on Radiation Protection and Measurements (NCRPM) requires that all x‐ray tubes operating above 70 kVp must include filtration that is equivalent to at least 2.5 mm aluminum. This applies to most veterinary practices. There is some inherent filtration in most x‐ray machines, but additional filters usually are needed.

Inherent filtration comes from the x‐ray tube glass envelop, the surrounding oil, the glass window in the protective housing, and the collimator. Inherent filtration typically contributes about 1.5 mm aluminum equivalent. Added filtration generally consists of aluminum plates that are positioned near the x‐ray tube in the path of the x‐rays (Figure 1.10). The plates provide the additional 1–2 mm of aluminum required.

In addition to lowering the radiation dose to the patient, filtration also increases the average or effective energy of the x‐ray beam. Removing the low energy x‐rays from the beam is called beam hardening. Beam hardening increases the effective energy of the x‐rays from about 1/3 of kVp to about 1/2 of kVp. A higher effective energy means the x‐ray beam is more uniform in intensity, which can improve radiographic quality.

Selective filtration sometimes is used to make the x‐ray beam more intense on one side than the other. This can be helpful when the thickness of the body part varies significantly. Rather than making two different exposures, one for the thick side and another for the thin side, a compensation filter can be used (Figure 1.11). A typical compensation filter is a wedge‐shaped piece of aluminum that unilaterally absorb x‐rays. It makes the x‐ray beam less intense on the thin side to prevent overexposing that side of the body part while allowing full exposure on the thicker side. The resulting radiograph is more uniform in density across both thicknesses. The heel effect provides a similar benefit and is discussed on page 16.

Schematic illustration of compensation filter. — **Figure 1.11** Compensation filter. In this illustration, the dog’s head varies in thickness from the nose to the ears. A wedge filter is used to reduce the amount of x‐rays to the nose while allowing the full intensity of the x‐ray beam to reach the thicker skull. This allows the head to be imaged without overexposing the nose or underexposing the skull. The gradual slope of the wedge blends the different x‐ray beam intensities together, so there is no sharp demarcation between the lighter and darker areas in the radiograph.

Compensation filters are attached to the front of the collimator, usually with magnets or Velcro. Some collimators are equipped with an adjustable holder designed to fit a compensation filter. The filter is installed after setting the FOV because the light from the collimator will be blocked by the filter. Compensation filters are less important with digital radiography because computer processing can be used to manipulate the image and adjust for variations in density.

Electrical supply to x‐ray machine

The x‐ray machine must receive a constant level of power to produce x‐rays in a consistent and reliable manner. Fluctuations in power can cause significant variations in x‐ray production. The line of incoming electricity should be dedicated to the x‐ray machine with no other equipment on the line. Appliances such as a clothes dryer or an air conditioner will pull power away from the x‐ray machine. A line voltage compensator is a useful device that adjusts for variations in incoming voltage. A voltage compensator is recommended because some power fluctuations are beyond the control of the hospital or clinic (e.g., originate at the power company).

The source of electricity for most x‐ray machines is alternating current (AC). Alternating current regularly reverses direction, switching the flow of electricity from positive to negative many times per second (Figure 1.12). Depending on the country of origin, AC electricity is either 50 or 60 Hz, which means it cycles from one direction to the other 50 or 60 times per second. Direct current (DC) moves in only one direction and does not reverse. X‐ray machines use both AC and DC power. However, the high voltage generator in the x‐ray machine converts AC to DC, so a separate DC supply is not needed. There are two different electrical circuits in an x‐ray machine, the filament circuit to produce electrons and the high voltage generator to accelerate the electrons.

Schematic illustration of high voltage generators. — **Figure 1.12** High voltage generators. This illustration depicts half‐wave and full‐wave rectification with single phase, three phase, and high frequency generators. The input power for single‐phase generators (A and B) is a single line of alternating current (AC). AC power oscillates the flow of electricity from a positive direction (+) to a negative direction (−) many times per second. X‐ray tubes require a current that moves in one direction only, which means the AC must be rectified to direct current (DC). Half‐wave rectifiers (A) remove the negative direction of the AC to produce pulses of DC electricity. During each pulse the voltage fluctuates from zero to the maximum and back to zero again, resulting in 100% ripple. Full‐wave rectifiers change the negative direction of AC to the positive direction to produce twice the number of DC pulses as half‐wave rectifiers. However, the voltage still fluctuates from zero to maximum with each pulse and the ripple is still 100%. Three‐phase generators (C) utilize three separate lines of input AC power. Each line is slightly out of phase with the other two. All three lines are full‐wave rectified to produce overlapping pulses of DC and a more continuous output of voltage than single‐phase units. As the voltage begins to drop in one pulse, the next pulse brings it back up to the maximum, resulting in only 15% ripple. High‐frequency generators (D) utilize either single‐phase or three‐phase AC input that is full‐wave rectified to DC. However, high‐frequency generators then “chop” the DC pulses into smaller and smaller pulses and “smooth” them to produce near‐constant output voltage with less than 1% ripple.

The filament circuit increases the amperage of the incoming AC power to provide enough current to heat the tungsten wire and produce a sufficient number of electrons. The amperage is increased using a step‐down transformer. A transformer is a device that increases or decreases voltage. Voltage and amperage together constitute power, and the power leaving a transformer must be the same as the power coming into it. The equation is W = V × A, where W is power measured in watts, V is voltage, and A is amperage. Any change in voltage must include an opposite change in amperage so the power can remain constant. The step‐down transformer, therefore, increases amperage to the filament by decreasing the incoming voltage.

The high voltage generator increases the voltage of the incoming AC power to provide the kilovolts needed to accelerate the electrons with enough speed to generate x‐rays. This is accomplished with a step‐up transformer. The step‐up transformer increases the incoming voltage 500–1000 times, which drops the amperage to only a few milliamps. This means the tube current (current across the x‐ray tube) is very low amperage and very high voltage.

The high voltage generator also converts AC to DC. X‐ray tubes are designed to use current that flows in one direction only: from the cathode to the anode. Current that flows in the other direction will not produce x‐rays and may damage or destroy the x‐ray tube. The process of converting AC to DC is called rectification. Rectification may be either half‐wave or full‐wave, both of which produce a DC output that consists of “pulses” of electricity (Figure 1.12).

Half‐wave rectifiers suppress the negative direction of AC, so only the electricity that flows in the positive direction is available. For example, if the input electricity is 60 Hz AC, the half‐wave rectifier would produce a DC output to the x‐ray tube of 60 pulses/second of electricity (Figure 1.12). As you can see, half‐wave rectifiers only use half of the input AC, the other half is wasted. Full‐wave rectifiers, on the other hand, change the negative direction of AC to the positive direction. Using the same example, a 60 Hz AC input would be converted to a DC output with 120 pulses/s, which is twice the electrical output of a half‐wave rectifier.

A sidenote that may be interesting: because x‐ray tubes only operate when current flows in one direction, a simple x‐ray tube that receives only AC power actually is self‐rectified. The problem with a self‐rectified tube is that the anode is likely to get hot enough to emit its own free electrons and then the current would flow in the reverse direction and destroy the cathode.

Because rectification produces DC output in pulses, the amount of electricity going to the x‐ray tube can fluctuate greatly. This fluctuation in power is called ripple (Figure 1.12). Specifically, ripple is the percent variation in voltage, from minimum to maximum, that is received by the x‐ray tube. The smaller the ripple, the more constant the voltage and the more uniform and intense the x‐ray beam. The amount of ripple depends on the type of high voltage generator. Currently, there are three basic types of generators: single‐phase, three‐phase, and high‐frequency.

Single‐phase generators receive a single source of AC power. The AC is rectified to produce a DC output with 100% ripple (Figure 1.12). 100% ripple means the x‐rays are produced in bursts instead of continuously. The intermittent production of x‐rays leads to longer exposure times, higher radiation doses, and a lower x‐ray beam intensity (only about 70% of peak output). The shortest exposure time with a single‐phase generator is 1/120 second (~8 ms).

Three‐phase generators receive three separate lines of AC power. Each line is superimposed on the other and 120° out of phase with the other two (Figure 1.12). All lines are full‐wave rectified so each line of 60 Hz AC produces 120 pulses of DC. The three lines are combined so together they produce 360 pulses/s of DC power. Because the lines are out of phase with each other, as the voltage in one pulse begins to drop another pulse brings it back up to the maximum so the voltage never falls to zero. The x‐ray tube receives more constant power with a ripple that is only about 15%. Some three‐phase generators include extra rectification to produce DC with 720 pulses/s and only 4% ripple. These are called three‐phase, 12‐pulse generators. Exposure times with three‐phase generators can be as short as 1/1000 second (1 ms), and the x‐ray beam intensity is about 95% of peak output.

High‐frequency generators supply near‐constant voltage to the x‐ray tube. The input power may be single‐phase or three‐phase AC which is then full‐wave rectified (Figure 1.12). High‐frequency generators then use high‐speed switches called inverters to effectively “chop” the DC pulses into smaller and smaller pulses. Inverters can produce DC with up to 25,000 pulses/s. All of these pulses are then “smoothed” using additional rectifiers and capacitors so the x‐ray tube receives voltage with less than 1% ripple. High‐frequency generators enable shorter exposure times, lower radiation doses, and help extend the life of the x‐ray tube. The x‐ray beam intensity may be over 99% of peak output. In addition, many high‐frequency generators are more compact, about 1/10 the size of a three‐phase generator.

Operating console

The x‐ray machine operating console includes controls to turn the equipment on and off and to adjust the mA, kVp, and the length of time the x‐rays are being produced.

The mA (milliamperes) control regulates the quantity of x‐rays produced. It determines the amount of current that is sent to the filament and therefore the number of electrons that are emitted. The higher the mA, the more electrons and the more x‐rays. The mA control has no effect on the energy or penetrating power of the x‐rays. Most x‐ray machines provide fixed settings for mA (i.e., 100 mA, 200 mA, 300 mA).

The kVp (kilovolt peak) control primarily regulates the energy or quality of the x‐rays. It regulates the amount of voltage that is sent to the x‐ray tube which determines the speed of the electrons. The higher the kVp, the greater the kinetic energy of the electrons and the greater the energy of the x‐rays.

The kVp control also contributes to the number of x‐rays produced. Higher kVp means electrons are more quickly removed from the cloud, which allows more electrons to be emitted from the filament. More electrons leads to more x‐rays.

The kVp control on most operating consoles displays continuously variable units of kilovoltage (kV). Voltage can be thought of as the “force” that moves the electrons from the cathode to the anode. Thousands of volts are required to accelerate the electrons with enough energy to create useful x‐rays. The maximum speed of the electrons is determined by the maximum or peak kilovoltage, which is called kVp. It is important to realize that kVp determines the peak energy of the x‐rays and not the total energy or the individual energy of the x‐rays. This is because not all of the electrons will be moving at maximum speed (due to the sine wave distribution of the voltage, as shown in Figure 1.12). The slower electrons produce less energetic the x‐rays. In addition, remember that many of the x‐rays produced at the anode will be less than maximum energy due to Bremsstrahlung. Therefore, the average or effective energy of most x‐ray beams is only about 1/3 of kVp. The effective energy can be increased to about 1/2 of kVp with high‐efficiency generators and beam hardening (filtration).

The exposure switch is a manual control that is used to start the production of x‐rays. The timer is an automatic control that stops x‐ray production. The exposure switch typically operates in two stages: a prep stage and an exposure stage. The prep stage prepares the x‐ray tube by rotating the anode and initiating the heating of the filament. It is important to make sure the anode is rotating at full speed before making an exposure, but it is equally important to avoid prolonged rotation as this will shorten the life of the x‐ray tube. During the exposure stage, x‐rays are being produced. The exposure stage is terminated automatically by the timer. There are a variety of timer mechanisms and most include a meter or digital display with units in fractions of a second (S) or milliseconds. In each case, the radiographer begins the exposure and the timer automatically stops it.

X‐ray table

The x‐ray machine table supports the patient during radiography. X‐ray tables are made of strong materials with low x‐ray attenuation, such as carbon fibers. Tables should be uniform in thickness and opacity and able to support at least 300 lb (136 kg). A slightly elevated rim or edge around the table top helps prevent liquids from spilling onto the floor or into the equipment under the table. Many x‐ray tables include a built‐in tray designed to hold a grid and various types and sizes of cassettes. The tray should be aligned with the x‐ray tube so the two can be moved simultaneously.

Floating tables help facilitate patient positioning. Most floating table tops can easily be locked, unlocked, and moved in multiple horizontal directions. The tray under a floating table generally remains aligned with the x‐ray tube while the tabletop is being moved. Tables that are height adjustable also facilitate patient positioning, especially given the variety of patient sizes commonly encountered in a veterinary practice.

Image receptors

Image receptors record the x‐rays that pass through an object or patient. The x‐rays received are used to form a latent image. The latent image is invisible and must be converted to a visible image for interpretation. Methods of conversion will be discussed shortly.

An image receptor must record all of the different x‐ray intensities that emerge from a patient and accurately translate them into acceptable levels of radiographic density (visible shades of gray). The minimum number of x‐rays a receptor can receive and still adequately darken the radiograph (produce a visible shade of gray) is known as its sensitivity. The more sensitive the system, the fewer x‐rays are needed. An image receptor must also display each level of density (shade of gray) with enough contrast so one can be distinguished from the next. This is called the dynamic range or latitude of the receptor. Wide latitude receptors are able to form many shades of gray over a large range of x‐ray intensities, which means they can display many different levels of opacity. Narrow latitude receptors form fewer shades of gray over a smaller range of x‐ray intensities. Narrow latitude means fewer levels of opacity can be displayed. The amount of radiographic detail an image receptor can provide determines how well we can discriminate the edges of structures that are close together.

Currently, there are two basic types of image receptors: film:screen and digital. Both types are used with conventional x‐ray machines.

Film:screen image receptors

The image receptors in traditional film radiography utilize special photographic film coupled with phosphorescent sheets called screens. The screens convert x‐rays into visible light. X‐ray film is much more sensitive to light than to x‐rays so the screens act to intensify the effect of the x‐rays on the film. This means fewer x‐rays are needed to adequately expose the film when using an intensifying screen. In a typical film radiograph, over 95% of the film exposure is due to light from an intensifying screen and less than 5% is from the x‐rays themselves. Intensifying screens enable lower x‐ray exposures, less radiation dose, and longer life of the x‐ray tube.

An intensifying screen consists of a thin polyester base coated on one side with phosphors. The phosphor layer is covered with a thin protective coating. The amount of light emitted by the phosphors is directly proportional to the amount of x‐rays received by the screen. Most phosphors are either a rare earth material that emits a green light or calcium tungstate which emits blue light. Modern x‐ray films are orthochromatic, which means the film is more sensitive to a specific color of light. It is important to match the phosphor emission color with the film absorption color.

X‐ray film consists of a thin, flexible, polyester sheet coated on one or both sides with an emulsion that contains silver halide crystals. The emulsion is covered with a thin protective layer. The plastic base is transparent and often tinted blue to reduce eye strain.

Silver halide crystals become structurally altered when struck by light from an intensifying screen or from direct interaction with an x‐ray. The altered crystals form the latent image. The latent image is then processed to become visible. Processing converts the altered silver halide crystals to black metallic silver. Film processing is discussed further below.

Both the intensifying screens and the x‐ray film commonly are enclosed in a protective, light‐proof container called a cassette (Figures 1.13 and 1.15). A cassette may contain one screen for use with single emulsion x‐ray film or two screens for double emulsion film. It is important to load the cassette with the correct type of film. Cassettes are available in a variety of sizes that correspond to different film sizes. Most cassettes swing open on a hinge, making it easy to remove and load the film. Screens usually are permanently mounted in the cassette, often on a flexible foam pad to maintain close, uniform contact with the film. The back of the cassette frequently is made of a radiation absorbing material to block backscatter radiation.

Schematic illustration of film:screen image receptor. — **Figure 1.13** Film:screen image receptor. In this picture, the cassette is open and standing on end with a sheet of x‐ray film in front. Intensifying screens are mounted inside the cassette, one on each side. The rectangular cutouts along the upper edge of each screen are for labeling the radiograph for identification (ID).

Cassettes can last for a long time with proper care. They should be regularly inspected for damaged hinges, loose screens, light leaks, etc. (Box 1.3). Do not store cassettes in prolonged heat as this will damage the intensifying screens. Numbering each cassette is useful to help identify problems that may arise. Use a permanent marker to write the number on an inconspicuous part of a screen, such as near the ID label or in a corner (Figure 1.14). Write the same number on the outside of the cassette. If an artifact appears in a radiograph, you’ll know which cassette to inspect for a possible cause.

Schematic illustration of numbered cassette. — **Figure 1.14** Numbered cassette. A. The number “6” is written on an intensifying screen (arrow) in the corner of an open cassette. B. In a radiograph made using this cassette, the number “6” is visible in the corner (arrow). Numbering the cassettes helps identify those with problems.

Film processing

The process of converting the invisible latent image into a visible radiograph involves immersing the film in different chemicals for specific periods of time. Film processing may be performed manually or automatically in a machine. With manual film processing, the chemical solutions are stored in dip tanks and the technician transfers the film from one tank to another, paying special attention to the time intervals (Box 1.4). With automatic processors, the film quickly moves from one solution to the next without human intervention. Automatic processors operate at higher temperatures to speed up the chemical reactions and shorten the processing time. The chemicals used in automatic processors differ from those used for manual processing in that they are specifically designed to work at high temperatures. Whether manual or automatic, film processing involves five basic steps:

Develop.
Rinse.
Fix.
Wash.
Dry.

Schematic illustration of film:screen system. — **Figure 1.15** Film:screen system. This illustration depicts a cross‐sectional view of a loaded film cassette. A sheet of x‐ray film (colored light blue) is sandwiched between two intensifying screens (colored green). The film is coated with emulsion on both sides. Three x‐rays, labeled A, B, and C, are depicted interacting with the film:screen system. **X‐ray A** strikes a phosphor in the top screen causing it to emit light in all directions. The light that hits the film alters some silver halide crystals in the upper film emulsion, which will produce a black area when the film is developed (depicted by the black oval). **X‐ray B** passes through the top screen without interacting with any phosphors and strikes the film directly. It alters fewer silver halide crystals than light from the screen (thus, a smaller black oval will be produced when the film is developed). **X‐ray C** passes through all of the upper layers to strike a phosphor in the bottom screen. The light emitted from the screen travels in all directions and alters some of the silver halide crystals in the bottom film emulsion. This is the advantage of double emulsion x‐ray film: it increases the likelihood of x‐ray detection so fewer x‐rays are needed to produce acceptable radiographic density. The backing in the cassette is designed to block backscatter x‐rays from exposing the film.

The developer is a chemical that converts the silver halide crystals to black metallic silver. The altered crystals are converted more quickly than the unaltered crystals, but given enough time the developer will eventually change all of the crystals to black. Therefore, the time the film spends in contact with the developer must be carefully controlled. Development is stopped by rinsing the film with water to dilute and remove the chemicals. Automatic processors incorporate rollers to squeeze the rinse water and chemicals away from the film.

The fixer is a chemical that dissolves any undeveloped silver halide crystals so they can be washed from the film. Fixer also “hardens” the film emulsion to help preserve the image. After an appropriate length of time in the fixer, the film is washed with water to remove the dissolved crystals and any chemicals. Automatic processors again use rollers to squeeze away the wash water and chemicals. The film is then dried for viewing and storage. Film radiographs should always receive a final viewing after they are dry, as the appearance of the radiograph may differ between wet and dry.

Film processing can significantly affect the quality of the radiograph. It is critical to maintain the developer and fixer at proper concentrations and temperatures per the manufacturer’s recommendations. Pay special attention to film immersion times and sequences. Be careful to avoid rough handling of film and splashing of chemicals. All processing equipment should be periodically inspected to ensure peak performance and proper chemical replenishment rates.

Disposal of processing chemicals must be in accordance with environmental protection rules and regulations. Companies and special equipment are available for reclaiming silver from fixer and washer solutions and from x‐ray film that is no longer needed. Reclaimed silver can be both environmentally and financially rewarding.

Film fog

Silver halide crystals can be altered by conditions other than light from intensifying screens or x‐ray interactions. These include exposure to ambient light, heat, pressure, chemicals, and aging. Crystals that are altered by conditions other than those caused by x‐rays add density to the radiograph, but do not provide any useful information. Unwanted blackening of x‐ray film is called fog. Fog makes it more difficult to identify structures in the image (Figure 1.16 and 1.36). Many of the causes of x‐ray film fog can be avoided if we know about them. See the Artifacts section in Chapter 2 for figures depicting film fog.

Schematic illustration of fog. Two photographs of a lakeshore depicting the appearance of structures with fog (A) and without fog (B). — **Figure 1.16** Fog. Two photographs of a lakeshore depicting the appearance of structures with fog (A) and without fog (B).

Light fog is caused by exposing unprocessed x‐ray film to visible light. Light fog can be avoided by storing unused film in light‐proof containers. Also, make sure the darkroom is light‐proof, and the safelight is the correct color and intensity (see safelight test in the Artifacts section of Chapter 2). Static electricity is a common cause of light fog and can be minimized by electrically grounding darkroom personnel, controlling the humidity, and carefully handling the x‐ray film.

Heat fog is caused by prolonged exposure of unprocessed x‐ray film to temperatures above 70 °F (20 °C). Avoid heat fog by storing unused film in a controlled environment.

Pressure fog is caused by physically compressing unprocessed x‐ray film. Pressure fog can result from storing boxes of unused film flat instead of upright. Over time the weight of the top sheets will fog the bottom sheets. Pressure fog also is caused by bending or creasing the film (Figure A23) or by sliding it across a table or countertop.

Chemical fog is caused by exposing unprocessed x‐ray film to the vapors from cleaning solutions, ammonia, formaldehyde, spare developer, and others. Avoid chemical fog by storing unused film away from these and similar agents and making sure your fingers are clean and dry before handling the film.

Aging fog occurs with expired x‐ray film. Over time, the unused film will “self‐process” and become darker. Aging fog can be avoided by paying attention to expiration dates and not using expired film for radiography.

Characteristics of film:screen systems

The thickness of the emulsions and the sizes of the grains in a film:screen image receptor determine its sensitivity and the amount of radiographic detail it can provide. Grains are the silver halide crystals in the film emulsion and the phosphors in the intensifying screens. Both sensitivity and detail are directly related to the sizes of the grains, but they are inversely related to each other. The larger the grains, the more sensitive the system, but the less detail it will provide. The smaller the grains, the better the detail, but more x‐rays will be needed to adequately darken the image.

Sensitivity in film radiography sometimes is called speed. Speed is a term from the early days of radiography when the x‐ray machine controls for mA and kVp were limited and setting the exposure time was the most variable factor. Film:screen systems that could accommodate short exposure times and still produce acceptable density in a radiograph were called high speed. Systems that required longer exposure times were called slow speed. The term “speed” has persisted even though modern x‐ray machines allow intricate mA and kVp manipulations in addition to exposure times.

The sensitivity and detail of a film:screen system is built‐in. The only way to change it is to switch to a system with different size grains or a different emulsion thickness. A film:screen system with larger grains can produce acceptable radiographic density with fewer x‐rays. When fewer x‐rays are needed, shorter exposure times are possible and larger body parts can be imaged. However, radiographic detail will suffer because larger grains are more visible in the film radiograph, resulting in a “mottled” or “grainy” appearance. A film:screen system with smaller grains produces better detail, however, higher x‐ray exposures will be needed to adequately darken the radiographs. High detail film:screen systems generally are limited to imaging smaller body parts.

Digital image receptors

Digital image receptors convert x‐rays into electronic signals. The signals are then digitized and sent to a computer. The computer transforms the digital image into an analog radiograph so it can be viewed and interpreted by humans. Newer technologies are constantly evolving and digital systems are continuing to improve. The currently available systems may be broadly classified as either CR (computed radiography) or DR (digital radiography). DR systems can be further divided into three general categories: (1) those that directly convert x‐rays into electronic signals, (2) those that indirectly convert x‐rays into electronic signals, and (3) CCD, which stands for charged coupling device. Whether CR or DR, all‐digital image receptors incorporate four basic steps:

Capture the x‐rays to form a latent image.
Translate the latent image into electronic signals.
Convert the signals to digital data for the computer.
Process the data to form an analog image.

Computed radiography

A computed radiography (CR) image receptor captures x‐rays using an imaging plate with photostimulable storage phosphors (PSP). The imaging plate is enclosed in a cassette that is similar in size, look, and feel to the cassettes used in film radiography (Figure 1.17). CR and film:screen sometimes are called “cassette radiography.”

Schematic illustration of CR image receptor. — **Figure 1.17** CR image receptor. This illustration depicts a cross‐sectional view of a CR cassette. The typical CR imaging plate is about 0.3 mm thick and composed of a stiff base material coated with a layer of photostimulable phosphors (PSP). The PSP grains are held in a polymer binding material. They absorb and temporarily store x‐ray energy. The surface of the imaging plate is protected by a tough plastic coating. The backing material between the base and the PSP layer is designed to block backscatter x‐rays.

When exposed to x‐rays, PSP temporarily stores captured x‐ray energy to form a latent image. After exposure, the CR cassette is manually transferred to a machine that removes the imaging plate and scans it with a laser beam (Figure 1.18). The laser light stimulates the PSP to release the stored x‐ray energy as light, a process known as photostimulation. The amount of light emitted is proportional to the captured x‐ray energy. A light guide collects the emitted light and sends it to a photomultiplier tube (PMT). The PMT converts the light into electronic signals, amplifies the signals, and sends them to an analog‐to‐digital converter (ADC). The ADC digitizes the signals and sends them to a computer for processing. While it is scanning the imaging plate, the laser removes most of the latent image. Any remaining latent image is erased by exposing the plate to bright light, making the cassette immediately ready for re‐use. With proper care, CR cassettes can be used thousands of times.

Schematic illustration of CR read-out. — **Figure 1.18** CR read‐out. This illustration depicts a CR imaging plate inside the CR reader. The plate is transported along rollers as a laser beam sweeps back and forth scanning the latent image. The red laser light stimulates the phosphors to release their stored energy as blue light. The blue light is captured in a light guide (LG) and channeled to a photomultiplier tube (PMT). The PMT converts the light into electronic signals and sends them to an analog‐to‐digital converter (ADC). The ADC digitizes the signals and sends them to a computer for processing and analog display.

CR imaging plates are always “on,” meaning they are always sensitive to nearby radiation. Nearby radiation includes that coming from the environment (background radiation) and scatter from radiographing other patients. CR cassettes should not be stored near the x‐ray machine and any cassettes that haven’t been used for several days should be placed in the CR reader and erased prior to use.

CR image receptors have been around since the mid‐1980s and closely emulate traditional film:screen systems. They are compatible with most existing x‐ray facilities and often utilize the same exposure techniques as for film radiography. CR cassettes are portable, enabling them to be used with multiple radiography facilities and different x‐ray machines as well as for positional radiography and both table‐top and under‐the‐table techniques. The initial cost of a CR system often is lower than many DR systems. Compared to DR, however, CR requires more user intervention, which means it takes longer to complete a radiographic study and patient throughput is slower.

Digital radiography

Digital radiography (DR) image receptors capture x‐rays using either a photoconductor or a scintillator. A photoconductor translates x‐ray energy directly into electrical charge as a one‐step process. A scintillator is a crystalline structure that converts x‐ray energy into visible light and then channels that light through thin crystals to a photodiode. The photodiode converts the light into electrical charge. This is a two‐step process that indirectly converts x‐ray energy into electrical charge. Whether formed directly or indirectly, the electrical charge is proportional to the energy of the x‐rays and forms the latent image.

Transistors are used to convert electrical charge into electrical signals. Each transistor contains a detector element that corresponds to a pixel in the digital radiograph. A group of thin‐film transistors called a TFT array produces the electronic signals, which are then digitized by an ADC and forwarded to a computer (Figure 1.19).

Schematic illustration of DR image receptor. — **Figure 1.19** DR image receptor. This illustration depicts the key components in a DR image receptor. The x‐ray converter either directly or indirectly translates x‐ray energy into an electrical charge. It is layered over a thin film transistor array which is mounted on a glass base. The TFT array converts electrical charge to electronic signals and sends them to an analog‐to‐digital converter, where they are digitized and sent to a computer.

A charged coupling device (CCD) is a type of indirect DR image receptor. A CCD uses a scintillator to convert x‐rays to light which is collected by an optical system. The optical system may consist of mirrors, lenses, prisms, or fiber optics. It captures the light and sends it to an array of CCD sensors. CCD sensors both convert the light into electrical charge and convert the electrical charge into electronic signals, without the need for photodiodes or a TFT array. CCD sensors are incorporated in many digital cameras.

As opposed to CR systems, DR image receptors convert captured x‐rays into electronic signals without the need for user intervention (Figure 1.20). In addition, DR systems are linked directly to a computer which further reduces the time needed to complete a radiography study. Compared to CR and film radiography, DR systems enable faster patient throughput and shorter turnaround times. DR image receptors can be permanently mounted under the x‐ray table, or they can remain portable for tabletop use and positional radiography (Figure 1.21).

Schematic illustration of types of DR image receptors. — **Figure 1.20** Types of DR image receptors. These illustrations highlight the different methods of converting x‐ray energy into electronic signals, which are then sent to be digitized and processed by a computer. **Direct DR systems** (A) capture x‐ray energy using a photoconductor which directly translate the energy into electrical charge (e⁻). The TFT array converts the electrical charge to electronic signals. **Indirect DR systems** (B) use a scintillator to convert captured x‐ray energy into visible light. The thin, needle‐shaped crystals in the scintillator channel the light to photodiodes, where it is converted into electrical charge (e⁻). Electrical charge is then converted to electronic signals by the TFT array. CCD (C) also uses a scintillator to convert captured x‐ray energy to light, but special sensors then convert both the light into electrical charge and the electrical charge into electronic signals.

Characteristics of digital systems

Unlike film:screen systems, digital image receptors do not have a fixed sensitivity. They are able to capture many more of the incident x‐rays and successfully convert them into acceptable radiographic density. How accurately and efficiently a digital receptor does this is described by various measurements including the modulator transfer function and the detective quantum efficiency, both of which are briefly discussed below.

Not all parts of a digital image receptor can convert x‐rays into electronic signals. This means that digital sensitivity is affected not only by the number of x‐rays that reach the receptor but also where they land. The area of the receptor that is sensitive to x‐rays in relation to the total area of the receptor is called the fill factor. The larger the fill factor, the fewer x‐rays will be needed to produce adequate density.

Modulator transfer function (MTF) measures how accurately a digital receptor transfers the different x‐ray intensities into different levels of radiographic density. The higher the MTF, the more different opacities can be displayed. MTF ranges from 0 (no structures in the patient are visible) to 1 (all structures are visible). An MTF of 0.5 means the receptor can display 50% of the differences in opacity as different radiographic densities.

Detective quantum efficiency (DQE) measures how efficiently a digital receptor captures the x‐rays it receives and converts them into acceptable radiographic density. The higher the DQE, the fewer x‐rays are needed. For example, a DQE of 0.4 means the receptor can utilize 40% of the x‐rays it receives. DQE is the ratio of input SNR to output SNR (SNR is signal‐to‐noise ratio). The higher the DQE, the better the radiographic detail.

The exposure index (EI) is a measure of the amount of x‐rays the digital image receptor actually received during the making of a particular radiograph. EI was standardized in 2018 and can be used to compare the sensitivities of different digital systems. EI is especially valuable to determine whether the exposure used was too high. In these cases, EI can be used to limit dose creep. Dose creep refers to the unnecessarily high radiation doses that frequently are encountered in digital radiography (dose creep is discussed later in this chapter).

EI is directly related to the x‐ray exposure dose. With most digital radiography systems, it should be in range of 1700–2100. If the exposure is doubled, the EI increases by 300. If the exposure is cut in half, the EI decreases by 300. Here’s how EI can be used: suppose a particular radiographic study resulted in an EI of 2500. An EI of 2500 is above the desired range of 1700–2100. This means the exposure used was too high and the radiograph should be overexposed. However, most digital systems can still display a good quality radiograph from an overexposure, so the image may appear acceptable. But the radiation dose to the patient and nearby personnel was too high. Overexposures need to be discussed with staff and prevented whenever possible. Any future images of this patient should be made using lower exposure. To bring an EI of 2500 into the desired range, we would cut the mAs in half (which reduces the EI by 300 to an EI of 2200) and cut the mAs in half again (which reduces the EI another 300 to an EI of 1900, which is in the desired range).

Sensitivity number (S‐number) is inversely proportional to x‐ray exposure. As exposure increases, the S‐number decreases. For most radiographic studies, the S‐number should be between 200 and 300. For example, a radiograph made with an S‐number of 100 is below the desired range of 200–300. This means the exposure used was too high (about twice what was needed). Future radiographs should be made with half the mAs, which would make the S‐number 200 and in the acceptable range. Another example: a radiograph made with an S‐number of 1200 is above the desired range of 200–300, which means the exposure was too low. Future exposures should be higher. Doubling the mAs reduces the S‐number to 600 and doubling the mAs again decreases the S‐number to 300, which is in the desired range.

The amount of radiographic contrast a digital receptor can display is determined by its pixels. “Pixel” is short for picture element. Pixels are the smallest pieces of a digital radiograph. Each pixel can appear as any of several different shades of gray, depending on its pixel number. Pixel numbers are binary numbers. Binary is a “base 2” numbering system that is used by computers, as opposed the “base 10” numbering system we humans use in our everyday lives (i.e., a decimal system). Each digit in a binary number is called a bit. “Bit” is short for b inary digit, and it is the smallest unit of digital data. Each bit is a digit in a binary number and is assigned a value of either 0 or 1. The position of each bit in the binary number represents an exponent of 2. In the decimal system, the position of each digit represents an exponent of 10.

In digital radiography, the contrast in an image can be adjusted after the radiograph has been acquired. This is a tremendous advantage over film radiography. Digital contrast can be optimized using the large amount of data in each pixel. Computer processing and user manipulations determine the best shades of gray for the pixels to display. However, for a computer to produce quality radiographs, it requires a properly calibrated image receptor and correct input from the user. The method of receptor calibration usually is built into each manufacturer’s system. Calibrations should be done per the manufacturer recommendations prior to radiography to correct any defects, imperfections, and non‐uniformities in the image receptor.

Geometry of the x‐ray beam

It is essential to understand the effects of x‐ray beam geometry and how to use them to make quality radiographs. The radiographic appearance of an object is greatly influenced by its position and orientation relative to the x‐ray beam and to the image receptor.

Focal spot

The focal spot is the source of x‐rays. It is the point on the anode target that is struck by the high‐speed stream of electrons. The focal spot, however, is not a tiny dot without length or width; it has dimensions. X‐rays are emitted from all parts of the focal spot and travel in straight lines in all directions. The x‐rays coming from one side of the focal spot overlap with those coming from the other side. This overlap makes the edges of objects appear blurry in the radiograph (Figure 1.22). Edge‐blurring is called geometric unsharpness. Geometric unsharpness reduces radiographic detail, making it more difficult to distinguish two objects that are close together. Edge‐blurring sometimes is called penumbra, which is an astronomy term used to describe the fuzzy‐edged shadow cast by the moon during a partial eclipse. The term umbra refers to the more sharply defined center part of the shadow.

Schematic illustration of geometric unsharpness. — **Figure 1.22** Geometric unsharpness. If x‐rays originated from a point source, as shown in illustration A, the margin of the object would appear sharp in the image below. However, x‐rays originate from a focal spot with dimensions, as shown in illustration B. X‐rays coming from one side of the focal spot (orange arrows) overlap with those coming from the other side (yellow arrows), which causes the object margin to appear blurred in the image below.

The degree of geometric unsharpness depends on the ratios of the distances between the focal spot, the object, and the image receptor and on the size of the focal spot. These ratios are explained in greater detail when we discuss distortion later in this chapter. The size of the focal spot may be altered by changing its physical size or its effective size. The smaller the focal spot, the less geometric unsharpness and the better the radiographic detail.

The physical size of a focal spot is fixed by the x‐ray tube manufacturer. Most x‐ray tubes are equipped with two focal spots; a small one and a large one (Figure 1.23). A small focal spot typically is about 0.5–1.0 mm in size and a large focal spot is about 1.0–2.0 mm. The small focal spot produces better radiographic detail than the large focal spot, but it is less tolerant of heat. Fewer and less energetic x‐rays can be produced at the small focal spot, which generally limits its use to body parts measuring less than 10–12 cm thick. The large focal spot is more heat tolerant and allows greater x‐ray exposures, shorter exposure times, and imaging of larger body parts, but there is less radiographic detail.

Schematic illustration of physical size of the focal spot. — **Figure 1.23** Physical size of the focal spot. These illustrations depict the effect of focal spot size on edge sharpness. The anode is shown in red at the top of each image. (A) With a small focal spot there is less overlap of the x‐rays and less blurring of the object margin in the image below. (B) With a larger the focal spot there is greater overlap of x‐rays and more geometric unsharpness.

The effective size of the focal spot can be altered using the line focus principle. The line focus principle states that viewing a sloped surface at an angle reduces its apparent size. This is why the anode disc is beveled. Tilting the anode target makes the focal spot appear smaller (Figure 1.24).

Schematic illustration of line focus principle. — **Figure 1.24** Line focus principle. These illustrations depict variations in the effective size of the focal spot due to the slope of the anode target and the size of the electron beam (e⁻). The steeper the slope, the smaller the effective size of the focal spot. The actual size of the focal spot is determined by the size of the electron beam. (A) 45° slope. (B) 20° slope. (C) The actual size of the focal spot is smaller because the stream of electrons is smaller (due to using a smaller filament). x = x‐rays.

The effective focal spot size also can be reduced by moving the x‐ray tube further away from the image receptor. With increased distance, the focal spot more closely resembles a point source. The greater the distance, the better the radiographic detail. However, as you increase the distance, the x‐ray exposure must be increased exponentially to maintain radiographic density. If you double the distance, you will need four times the x‐ray exposure due to the inverse square law.

Inverse square law

The inverse square law occurs because x‐rays diverge as they move away from the focal spot, spreading out over larger and larger areas (Figure 1.25). This means that there are fewer x‐rays per unit area (less x‐ray beam intensity) as you move further away from the source. The decrease in x‐ray beam intensity is inversely proportional to the square of the distance from the focal spot: I = 1/D ², where “I” is the intensity of the x‐ray beam and “D” is the distance from the focal spot.

Heel effect

Because the target is angled to create a smaller effective focal spot, the bottom of the anode is thicker. The thicker bottom commonly is called the “heel” of the anode. The heel absorbs some of the x‐rays produced at the focal spot. Absorption results in fewer x‐rays being emitted from the anode side of the x‐ray tube and more x‐rays coming from the cathode side. The heel effect can produce as much as a 40% difference in x‐ray beam intensity between the anode side and the cathode side of the x‐ray tube (Figure 1.26). This difference is useful when imaging body parts that vary in thickness or opacity. The thicker or more opaque end of the body part is positioned toward the cathode side where the x‐ray beam is more penetrating. Remember, the word “cathode” has more letters than the word “anode” and more x‐rays come from the cathode side than the anode side. The heel effect is most noticeable when using a:

Large FOV.
Lower x‐ray exposure.
Short distance between the x‐ray tube and the image receptor.

Schematic illustration of heel effect. — **Figure 1.26** Heel effect. More x‐rays are emitted from the cathode side of the x‐ray tube than from the anode side because some of the x‐rays produced are absorbed in the anode heel. There is a gradual loss of x‐ray beam intensity moving from cathode side to anode side.

Distortion

Distortion occurs when an object’s actual size or shape is misrepresented in the radiograph. It is caused by equal or unequal magnification of the object. Equal magnification makes an object appear overall larger in the radiograph than its actual size. Unequal magnification causes foreshortening or elongation distortion which are described below. Any form of magnification also causes geometric unsharpness. There are three distances that are important in distortion:

Distance between the focal spot and the image receptor: the source‐to‐image distance or SID.
Distance between the focal spot and the object: the source‐to‐object distance or SOD.
Distance between the object and the image receptor, the object‐to‐image distance or OID.

It is the ratio between any two of these distances that is important when dealing with magnification and geometric unsharpness (Figure 1.27). To minimize magnification, we want the object to be as close to the image receptor as possible (OID as short as practical). The degree of magnification is determined by the ratio of SID to SOD (SID/SOD is called the magnification factor). The closer the ratio is to 1.0, the less magnification (Figure 1.28).

Schematic illustration of magnification. — **Figure 1.27** Magnification. (A) Source‐to‐image distance (SID), source‐to‐object distance (SOD), and object‐to‐image distance (OID) are illustrated. (B) Increasing the OID (moving the object further away from the image receptor) while maintaining the same SID creates magnification distortion and geometric unsharpness. (C) Shortening the SID (moving the x‐ray tube closer to the image receptor) while maintaining the same OID also creates magnification. (D) Shortening the OID (placing the object closer to the image receptor, such as using a table‐top‐technique instead of under the table) and using the same SID will lessen magnification.

Schematic illustration of magnification factor. — **Figure 1.28** Magnification factor. Illustration A depicts an object close to the image receptor with little magnification. The SID and SOD are similar resulting in a magnification factor (MF) near 1.0. In illustration B, the object is further from the image receptor, resulting in magnification and geometric unsharpness. The object appears twice as large in the image as actual size (magnification factor is 2.0).

Sometimes magnification is desired. Here is an example: suppose we want to make a radiograph that displays the magnification of a paw. We position the patient’s foot on a cardboard box or foam block to move it further away from the image receptor, increasing the OID. This will cause magnification and geometric unsharpness. To minimize the edge‐blurring, we use the small focal spot and move the x‐ray tube further away, increasing the SID. Because the SID is longer, the x‐ray exposure needs to be increased due to the inverse square law, so we increase the mAs.

Both elongation distortion and foreshortening distortion occur when the position of an object is oblique to the x‐ray beam or image receptor. Oblique positioning makes the OID longer on one side of the object than the other. When one side is magnified more than the other the shape of the object is distorted. Unequal magnification also can occur if the image receptor is obliquely positioned in relation to the object or x‐ray beam. Foreshortening distortion makes an object appear overall shorter in the radiograph than its actual size (Figure 1.29). Elongation distortion makes an object appear overall longer than its actual size (Figures 1.30 and 1.31).

Schematic illustration of foreshortening distortion. — **Figure 1.29** Foreshortening distortion. A. The long axis of the bone is parallel with the image receptor and perpendicular to the x‐ray beam and therefore its size and shape are accurately represented in the image below. B. The bone is tilted in relation to the image receptor and oblique to the x‐ray beam so that it appears shorter than actual and abnormal in shape due to unequal magnification.

Schematic illustration of elongation distortion. — **Figure 1.30** Elongation distortion. Object A is aligned with the central ray (red arrow) resulting in an accurate depiction of its size and shape in the image below. Object B is off‐center, peripheral to the central ray, making it appear longer and more ovoid than actual in the image below.

Elongation also occurs in objects that are off‐center to the x‐ray beam. Because x‐rays diverge, objects nearest the center of the beam will be imaged more accurately than objects at the periphery of the beam. All peripheral objects in a radiograph suffer some degree of elongation distortion. For example, when imaging the spine only the vertebrae and intervertebral disc spaces in the center of the image are accurately depicted for size and shape. The peripheral vertebrae are more elongated than actual, which may be mistaken for narrowed intervertebral disc spaces. Another example is the appearance of the cardiac silhouette at the periphery of an abdominal radiograph. Due to elongation distortion the cardiac silhouette is distorted and may appear longer than actual, which may be mistaken for cardiomegaly.

To minimize distortion, position the object of interest as close to the image receptor as practical (minimize the OID), align the object with the center of the x‐ray beam, and collimate the FOV as small as practical. When interpreting radiographic findings, be aware of distortion artifacts.

X‐ray interactions with matter

Each x‐ray that enters a material will result in one of three possible outcomes: (1) it may be transmitted, passing through the material unchanged, (2) it may be absorbed and completely disappear inside the material, or (3) it may be scattered, becoming a secondary x‐ray that travels in a new direction with less energy (Figure 1.32). Transmitted and absorbed x‐rays are valuable to make radiographs. Scatter x‐rays provide no useful diagnostic information, degrade image quality, and expose the patient and nearby personnel to unnecessary radiation.

Schematic illustration of x-ray interactions. — **Figure 1.32** X‐ray interactions. X‐rays produced in the x‐ray tube are called *primary* x‐rays. Primary x‐rays that pass through the patient unchanged are *transmitted* x‐rays. Primary x‐rays that interact with atoms in the patient may either be *absorbed* (lose all of their energy and completely disappear) or be *scattered* (become a secondary x‐ray that travels in a new and unpredictable direction, as shown by the red arrows).

Transmitted x‐rays that reach the image receptor and are recorded produce useful blackened areas in a radiograph (i.e., radiographic density). The more transmitted x‐rays, the darker that part of the image. The likelihood of an x‐ray being transmitted is affected by its energy. The higher the energy, the more penetrating the x‐ray. Recall that x‐ray energy is controlled by kVp, so the higher the kVp, the more transmitted x‐rays and the darker the radiograph. The likelihood of x‐ray transmission also is affected by the opacity of the material. The greater the opacity, the fewer x‐rays will pass through. Higher opacity materials appear lighter in radiographs. The mAs also will affect the number of transmitted x‐rays because mAs affects the overall quantity of x‐rays produced. The higher the mAs, the greater the number of transmitted x‐rays.

Absorbed x‐rays never reach the image receptor and therefore produce no radiographic density. The more absorbed x‐rays, the lighter or brighter that part of the image. In medical radiography, x‐ray absorption primarily occurs via the photoelectric effect.

The photoelectric effect is similar to characteristic x‐ray production (Figure 1.5) except it is an x‐ray instead of a high‐speed electron that collides with an atom. The x‐ray transfers all of its energy to an inner level electron, ejecting the electron and ionizing the atom (Figure 1.33). The ejected electron travels a short distance before losing its energy to other atoms along its path. Both the ionization of the atom and the energy from the ejected electron can damage or destroy living cells. Ionized atoms are unstable so a higher energy outer level electron drops down to fill the void in the inner level. The outer level electron loses energy in the form of an x‐ray. This new x‐ray is a secondary x‐ray because it was not produced in the x‐ray tube (primary x‐rays are emitted from the tube). The energy of the secondary x‐ray is equivalent to the difference between the outer level and inner level binding energies. In most cases, the secondary x‐ray is absorbed without further damage to cells.

Schematic illustration of photoelectric absorption. — **Figure 1.33** Photoelectric absorption. 1. A primary x‐ray collides with an atom in the patient. The x‐ray transfers all of its energy to an inner shell electron and completely disappears. 2. The electron now has enough energy to escape the atom and is ejected. 3. An outer shell electron drops down to fill the void in the inner shell. 4. The outer shell electron emits excess energy as a new x‐ray. The energy of the secondary x‐ray is less than that of the primary x‐ray (note the longer wavelength). N⁺ = nucleus; e^– = electrons.

The likelihood of x‐ray absorption via the photoelectric effect depends on the size of the atoms (i.e., atomic number) and the energy of the x‐rays (i.e., kVp). The larger the atomic number, the greater the chance of x‐ray absorption. Doubling the atomic number increases the probability of x‐ray absorption by a factor of 8. The higher the kVp, the less the chance of x‐ray absorption because the x‐rays are more penetrating. Doubling the x‐ray energy decreases the probability of absorption by a factor of 8.

The photoelectric effect is a major cause of x‐ray attenuation up to about 70 kVp. Above 70 kVp the probability of x‐ray absorption decreases, except at the K‐edge. At the K‐edge there is a sudden increase in absorption. K‐edge refers to the binding energy of an inner level electron (the inner level of an atom is called the K‐shell). X‐rays with energy that is slightly greater than the binding energy of an inner level electron (just above the K‐edge) are more likely to be absorbed (Figure 1.34).

The K‐edge can be quite useful when imaging materials with a high atomic number, particularly positive contrast media (e.g., barium, iodine). Materials with high atomic numbers contain more K‐shell electrons and therefore are more likely to absorb x‐rays with a certain energy. The K‐edge is not nearly as useful in soft tissues because these materials are composed of elements with smaller atomic numbers and fewer K‐shell electrons (e.g., carbon, hydrogen, oxygen, nitrogen). The K‐edges of barium and iodine are around 35 keV which is about the effective energy of a 70 kVp x‐ray beam. This is one reason why barium and iodine are such ideal positive contrast agents for medical radiography. When performing a positive contrast study try to keep the kVp as close to 70 as practical to enhance x‐ray absorption. The greater the x‐ray absorption, the greater the contrast.

Schematic illustration of k-edge absorption. — **Figure 1.34** K‐edge absorption. This graph depicts the relative abilities of fat, soft tissue, bone, barium, and iodine to absorb or attenuate x‐rays. The differences in attenuation are highest with lower x‐ray energies (lower kV). Radiographic contrast between the different tissues is greatest with lower energy x‐rays. As x‐ray energy increases, the x‐rays are more penetrating and less likely to be absorbed. There is less contrast between tissues when using higher energy x‐rays. However, at the K‐edges of iodine, barium, and to a lesser extent bone, there is a sudden increase in x‐ray absorption. At these parts of the graph there is a significant difference in attenuation between tissues which can be used to maximize radiographic contrast.

Scatter x‐rays are secondary radiation. During most radiographic procedures, about 95% of scatter comes from the patient and the rest from the x‐ray table, image receptor, floor, or any other material that happens to be in the x‐ray beam FOV.

Scatter x‐rays result from the incoherent or coherent interactions of primary x‐rays. Most scatter is caused by incoherent scattering, which is also called the Compton effect, inelastic scattering, or modified scattering. Incoherent scattering occurs throughout the range of x‐ray energies used in medical imaging (40–125 kVp). An x‐ray collides with an atom, ejecting an electron and ionizing the atom (Figure 1.35). The x‐ray energy is transferred to the ejected electron and to a secondary x‐ray. The secondary x‐ray travels in a completely random direction, anywhere from 0° to 180° from the original x‐ray’s path. Most incoherent scattering tends to be in the forward direction, but it is wholly unpredictable. If the secondary x‐ray travels back toward the x‐ray tube it is called backscatter radiation.

Tags: Radiography of the Dog and Cat Guide to Making and Interpreting Radiographs

Mar 16, 2025 | Posted by admin in ANIMAL RADIOLOGY | Comments Off

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