An x-ray is a discrete bundle of electromagnetic energy called a photon. In that regard, it is similar to other forms of electromagnetic energy such as light, infrared, ultraviolet, radio waves, or gamma rays. The associated electromagnetic energy can be thought of as oscillating electric and magnetic fields propagating through space at the speed of light. The various forms of electromagnetic energy differ only in frequency (or wavelength). However, because the energy carried by each photon is proportional to the frequency (the proportionality constant is called Planck's constant), the higher frequency x-ray or gamma ray photons are much more energetic than, for example, light photons and can readily ionize the atoms in materials on which they impinge. The energy of a light photon is of the order of one electron-volt (eV), whereas the average energy of an x-ray photon in a diagnostic x-ray beam is on the order of 30 kiloelectron volts (keV) and its wavelength is smaller than the diameter of an atom (10−8 cm).
In summary, an x-ray beam can be thought of as a swarm of photons traveling at the speed of light, each photon representing a bundle of electromagnetic energy.
Electromagnetic radiation may be produced in a variety of ways. One method is the acceleration or deceleration of electrons. For example, a radio transmitter is merely a source of high-frequency alternating current that causes electrons in an antenna wire to which it is connected to oscillate (accelerate and decelerate), thereby producing radio waves (photons) at the transmitter frequency. In an x-ray tube, electrons boiled off from a hot filament (Figure 2-1) are accelerated toward a tungsten anode by a high voltage on the order of 100 kilovolts (kV). Just before hitting the anode, the electrons will have a kinetic energy in kiloelectron volts equal in magnitude to the kilovoltage (eg, if the voltage across the x-ray tube is 100 kV, the electron energy is 100 keV). When the electrons smash into the tungsten anode, most of them hit other electrons, and their energy is dissipated in the form of heat. In fact, the anode may become white-hot during an x-ray exposure, which is one reason for choosing an anode made of tungsten, with a very high melting point. The electrons penetrate the anode to a depth less than 0.1 mm.
A small fraction of the electrons, however, may have a close encounter with a tungsten nucleus, which, because of its large positive charge, exerts a large attractive force on the electron, giving the electron a hard jerk (acceleration) of sufficient magnitude to produce an x-ray photon. The energy of the x-ray photon, which is derived from the energy of the incident electron, depends on the magnitude of the acceleration imparted to the electron. The magnitude of the acceleration, in turn, depends on how closely the electron passes by the nucleus. If one imagines a target consisting of a series of concentric circles, such as a dart board, with the bull's-eye centered on the nucleus, more electrons clearly will impinge at larger distances than in the bull's-eye; hence, a variety of x-ray photon energies will be produced at a given tube voltage (kV) up to a maximum equal to the tube voltage (a hit in the bull's-eye), where the electron gives up all its energy to the x-ray photon. Increasing the voltage will shift the x-ray photon spectrum to higher energies, and higher-energy photons are more penetrating. The radiation produced in this manner is called Bremsstrahlung (braking radiation) and represents only about 1% of the electron energy dumped into the anode by the electron beam; the other 99% goes into heat.
The electron current from filament to anode in the x-ray tube is called the mA, because it is measured in milliamperes. The mA is simply a measure of the number of electrons per second making the trip across the x-ray tube from filament to anode. The rate of x-ray production (number of x-rays produced per second) is proportional to the product of milliamperage and kilovoltage squared. The quantity of x-rays produced in an exposure of duration s (in seconds) is proportional to the product of mA and time and is called the mAs. The quantity of x-rays at a given point is generally measured in terms of the amount of ionization per cubic centimeter of air produced at that point by the x-rays and is measured in roentgens (R) or in coulombs per kilogram of air. This quantity is called exposure, and 1 R of exposure results in 2 × 109 ionizations per cubic centimeter of air.
The electron beam is made to impinge on a small area on the anode of the order of 1 mm in diameter in order to approximate a point source of x-rays. Because a radiograph is a shadow picture, the smaller the focal spot, the sharper the image. By analogy, a shadow picture on the wall (such as a rabbit made with one's hand) will be much sharper if a point source of light such as a candle is used rather than an extended light source such as a fluorescent tube. The penumbra (or unsharpness) of the shadow will depend not only on the source size, but also on the magnification, as can be illustrated by making a shadow of one's hand on a piece of paper using a small light source such as a single light bulb. The closer you bring your hand to the paper (the smaller the magnification), the sharper the edges of the shadow. Similarly, magnification of the x-ray image produced by the point source is less, the closer the patient is to the film and the farther the source is from the film. The magnification factor (M) is defined as the ratio of image size to object size and is equal to the ratio of the focal-to-film distance divided by the focal-to-object distance (M ≥ 1, and M = 1 means no magnification is produced; ie, either the object is right against the film, or the focal spot is infinitely far away). The penumbra, blurring, or unsharpness (Δx) produced on an otherwise perfectly sharp edge of an object and due to the finite focal spot size of dimension a is expressed by the equation
Unfortunately, the smaller the focal spot, the more likely it is that the anode will melt. The power (energy/per second) dumped into the anode is equal to the product of the kilovoltage and milliamperage; ie, at 100 kV and 500 mA, 50,000 watts of heat energy is deposited into an area on the order of a few square millimeters (imagine a 50,000-watt light bulb to get an idea of the heat generated).
Interaction of X-Rays with Matter
X-rays primarily interact with matter through interaction of their oscillating electric field with the atomic electrons in the material. Having no electrical charge, the x-rays are more penetrating than other types of ionizing radiation (such as alpha or beta particles) and are therefore useful for imaging the human body. The x-rays may be absorbed or scattered by the atomic electrons. In the absorption process (photoelectric absorption), the x-ray is completely absorbed, giving all its energy to an inner shell atomic electron, which is then ejected from the atom and goes on to ionize other atoms in the immediate vicinity of the initial interaction. In the scattering process (Compton scattering), the x-ray ricochets off an atomic electron, losing some of its energy and changing its direction. The recoiling electron also goes on to ionize hundreds of atoms in the vicinity. Electrons from both processes go on to ionize many other atoms and are responsible for the biological damage produced by x-rays.
The attenuation of the x-ray intensity with thickness of material follows an exponential law due to the random hit-or-miss nature of the interaction. The process is similar to firing a volley of rifle bullets into a forest, where the bullets may either stick in a tree (be absorbed) or ricochet off a tree (scatter). The deeper you go into the forest, the fewer bullets there are; however, a bullet still has a chance of traveling through the forest without hitting a tree. Likewise, an x-ray can make it all the way through a patient's body without touching anything and remain unchanged, as if it had passed through a vacuum instead. These are called primary x-rays. Typically, only about 1% of the incident x-rays penetrate the patient, and only about a third of these are primary x-rays; the rest are scattered x-rays that do not contribute to the anatomic image. An x-ray image is a shadow or projection image which assumes that x-rays reaching the film have traveled in a straight line from the source, but this is true only for the primary x-rays. As Figure 2-2 A shows, the film density (blackness) at point P on the film is related to the anatomy along line FP. The scattered photon reaches the film along the path FSP and is relaying information about the anatomy at the random point S to point P on the film. Scatter simply produces a uniform gray background; it does not contribute to the image. Because scatter reduces image contrast, it is desirable that the scatter be removed. This task is accomplished by use of an antiscatter grid (Figure 2-2 B). This grid consists of a series of narrow lead strips with radiolucent (low-attenuation) interspace material to remove some of the scatter. With the grid, the scattered photon shown in the figure can no longer reach the film, but the primary x-rays can. More of the scatter than primary x-rays is eliminated by the grid; hence, image contrast increases, but at the cost of an increase of a factor of 2 to 3 in patient dose. This increase occurs because the scatter, which was previously blackening the film, has been reduced, and therefore, higher x-ray exposure to the front of the patient is necessary to get the requisite number of x-rays through the grid to blacken the film. The grid is usually made to move a few interspaces during the exposure by a motor drive, in order to wash out the grid lines.
(A) Scattered and primary x-ray photons reaching the same point P on film. (B) Scattered photon is removed by antiscatter grid, while primary photon gets through.
The absorption process is more prevalent at lower kilovoltages and in materials with higher atomic numbers. Bones appear white on an x-ray film because photoelectric absorption of x-rays is greater in bone than in soft tissue as a result of the higher atomic number of bone. Lead is a useful shielding material for x-rays because of its high atomic number. The probability of the absorption process decreases rapidly with photon energy (as 1/E3) and the scattering process decreases slowly (as 1/E); hence, the x-ray beam becomes more penetrating as kilovoltage increases. The scattering process is roughly independent of the atomic number of the attenuating material (all electrons look alike to the photon for the scattering process), whereas the absorption process is more probable for tightly bound electrons such as the inner electrons in heavier elements.
Increasing the kilovoltage is therefore beneficial to the patient in that it reduces the radiation dose: that is, fewer x-rays must penetrate the front of the patient to get the requisite number out the back to blacken the film. However, an increase in kilovoltage will reduce image contrast because the absorption process, which is sensitive to atomic number, will decrease and the scattering process is independent of the atomic number of the materials. Even with materials of the same atomic number, contrast improves at lower kilovoltage settings because of higher attenuation, which results in greater differential attenuation between different thicknesses of the same material. Thus there is a tradeoff between image quality (contrast) and patient dose that must be weighed in the selection of kilovoltage.
For production of radiographic images, the x-ray film is placed in a cassette and sandwiched between two fluorescent screens that glow under x-ray exposure, and it is primarily the light from these fluorescent screens that blackens the film. Although x-ray film, which is quite similar to ordinary photographic film, can be blackened by direct x-ray exposure, the film does not absorb the penetrating x-rays very efficiently, because the emulsion consists of silver halide crystals embedded in a low-atomic-number gelatin base. The fluorescent screens, called intensifying screens, are made of high-atomic-number materials, which therefore absorb x-rays very efficiently and also emit hundreds of light photons per x-ray absorbed. These light photons, in turn, are efficiently absorbed by the film. As a result, x-ray exposure to the patient is reduced by a factor on the order of 100 compared to direct x-ray exposure of the film. The screens do produce a loss of sharpness of the image due to the spreading out of the light from the point of x-ray absorption before the light reaches the film. This effect can be reduced by making the screen thinner; however, it then absorbs a smaller fraction of the incident x-rays and therefore results in a “slower” system (more patient exposure is required).
In recent years digital image receptors have come into use. One type called CR (computed radiography) utilizes a cassette with a photostimulable phosphor material that stores the x-ray image in the form of trapped electrons for later readout by a scanned laser beam, which releases the electrons from their traps. On release, these electrons cause the phosphor to emit light that has a shorter wavelength than that of the laser beam. This light signal is read out and digitized, thereby forming a digital image. Another type called DR (direct radiography) consists of a flat-panel digital detector plate that is built into the x-ray unit itself. In these, the x-ray image is converted to an electrical signal from a fine matrix of thin-film transistor elements, which creates a digital image having a pixel size of 0.2 mm or less. These digital images, which consist of an array of numbers in a matrix, can be processed to improve image quality; displayed and manipulated on a viewing monitor; and then printed onto film using a laser film printer. The advantage of these digital systems is that the image can be processed to improve contrast and provide edge enhancement, and the film can be printed to the appropriate darkness regardless of the x-ray exposure.
Recall that the quantity of x-rays produced during an exposure is proportional to
However, because the beam is more penetrating at high kilovoltage, the x-ray exposure that reaches the film through a patient is roughly proportional to
That is, it depends very strongly on kilovoltage. The exposure time required to blacken the film is thus proportional to
The heat deposited in the anode is proportional to the product of kV and mAs.
Choice of an exposure technique is generally made by first selecting the kilovoltage. A lower kilovoltage gives greater image contrast but also higher patient exposure and requires a longer exposure time at a given milliampere setting, because the x-ray beam is less penetrating and x-ray production is lower at the lower kilovoltages. Thus, for thick body parts, care must be taken not to choose too low a kilovoltage.
Generally, x-ray tubes have two focal spot sizes produced by two different (selectable) filament sizes. That is, they have a large and a small focal spot (eg, 1.25 and 0.6 mm). With the small focal spot, however, the electron energy is deposited in a smaller area, thereby creating a higher anode temperature; hence, at a given kilovoltage, the maximum milliamperage that can be used without melting the anode is limited to a lower value, thereby resulting in a longer exposure time. The small focal spot will result in a sharper image, however, if the longer exposure time required by its selection does not “stop” patient motion; then motion of the patient during the exposure may blur out any sharpness gain realized by use of the small focal spot. In any case, the small focal spot is useful only for looking at fine detail, such as bony detail, and its use does not significantly improve, for instance, an abdominal radiograph in which soft-tissue contrast is the objective. The small focal spot might be used for radiographs of the skull or extremities. The exposure time selected should be short enough to stop the motion of the anatomic part being radiographed. A very short time would be required for the heart and somewhat longer times for the abdomen or chest. Exposure time is less critical for the head or extremities, which are not subject to motion in most cases.
Having selected the kilovoltage and exposure time, one must then select the milliamperage so that the milliampere-seconds (the product of milliamperage and time) is large enough to blacken the film suitably. If the milliamperage required is above 200 mA to 300 mA, a small focal spot generally cannot be used, because it will not allow this high a value of milliamperage without melting the anode.
On many x-ray units, a phototimer sensor (automatic exposure control) is used to automatically terminate the exposure when a given x-ray exposure has been accumulated at the cassette position. In this way, the film is blackened sufficiently regardless of patient thickness and kilovoltage selection. When using this feature, however, the operator loses control of the exposure time. Choosing the highest milliamperage allowable by the tube will ensure the minimum exposure time.
If, instead of using the light from a fluorescent screen to blacken a film, one viewed the fluorescent screen directly with the naked eye, then one would be performing fluoroscopy as it was done in the early days of medical x-ray use. Unfortunately, the image made in this fashion was very dim, even at a high exposure rate to the patient, so modern fluoroscopy uses an image intensifier that amplifies the light from a fluorescent screen. A typical fluoroscopic imaging system is shown in Figure 2-3. The image intensifier tube is an evacuated glass or metal tube with a fluorescent screen (input phosphor) that glows with the image produced by the x-ray pattern that exits the patient. The light from the input phosphor causes ejection of electrons from a photoelectric material adjacent to the input phosphor. These electrons are accelerated via a high voltage (30 kV), as well as being focused to preserve the image onto a small (1-inch diameter) screen (the output phosphor), which glows with the image because of the energy deposited by the impact of the accelerated electrons. The output phosphor glows much more brightly than the input phosphor (about 3000 times) because of the energy gain provided by the acceleration of the electrons and also because of minification of the image on the output phosphor. The image on the output phosphor can be viewed with the naked eye, usually with a series of lenses and mirrors, but the image is more commonly viewed by focusing a video camera onto the output phosphor and viewing the image on a TV monitor via a closed-circuit TV system. The fluoroscopic image generally has less contrast and less resolution of fine detail than a radiographic image; however, it is clearly convenient to view the image in real time—particularly when observing the flow of radiopaque contrast agents ingested or injected into the body. (These contrast materials, such as iodine or barium compounds, have a higher atomic number than soft tissue, hence, absorb more x-rays.) During fluoroscopic examinations, the x-ray tube is typically operated below 100 kV and below 3 mA tube current. Even so, entrance exposure rates (at the point where the x-ray beam enters the patient) are about 2 to 5 R/min, depending on patient thickness; hence, fluoroscopic examinations generally result in significantly higher patient exposures than do radiographic examinations.
Fluoroscopic imaging system.
Fluoroscopic systems generally have an automatic brightness control in which the brightness of the output phosphor is sensed by a light detector. The brightness signal from this detector is compared to a reference level, and the difference signal is used to instruct the x-ray generator to vary milliamperage or kilovoltage (or both) in order to maintain a constant brightness at the output phosphor. For example, after ingestion of barium in a barium-swallow examination, the barium absorbs significantly more x-rays, and the image would tend to go dark without such a system; however, as the brightness falls below the reference level, the automatic brightness control causes the x-ray generator to increase the milliamperage or kilovoltage to maintain a constant brightness on the monitor.
Recording of Fluoroscopic Images
Fluoroscopic images can be recorded for later viewing by several methods. The TV image can be recorded using a videotape recorder or a videodisc recorder, the latter having the advantage of being able to view one frame at a time as well as providing random access rather than the sequential viewing required by videotape.
In addition, some systems have the capability of digitizing the electric signal from a TV frame and storing it in computer memory chips. These systems often have a “last image hold” capability that holds the last TV frame on the monitor. This method is also used in digital subtraction angiography (DSA); that is, the analog signal from the TV camera is digitized and stored frame by frame in a computer memory in a 512 × 512 or 1024 × 1024 image matrix. A short radiographic x-ray pulse is usually used for making the image. Images made just before and after injection of contrast material into the arteries can be subtracted digitally, so that only the vascular system appears in the subtracted image.
The aforementioned image recording methods merely store the image recorded by the TV camera, which is of lower quality than a radiographic image and has even poorer resolution than the image appearing on the output phosphor of the image-intensifier tube because of the limitations of the TV imaging process. In order to record higher quality images during a fluoroscopic examination, spot film devices are used. The most common device transports a conventional radiographic screen/film cassette to a position in front of the image intensifier at the push of a button on the fluoroscopic carriage. The x-ray tube is then switched into a radiographic mode (ie, the milliamperage is increased from low mA to 200 to 400 mA to shorten exposure time), and a conventional radiographic image is obtained on film. Digital spot films can be obtained by digitizing a TV frame from the image intensifier acquired with a short exposure burst at a higher exposure value than that of a single continuous fluoroscopic frame. These produce an image of higher quality (lower noise) than that obtained from the fluoroscopic image.
In radiography or fluoroscopy, one is creating a shadow picture or a projection of the attenuation properties of the human body onto a plane. Thus, each ray from the source to a given point on the film, such as ray FP in Figure 2-2, conveys information about the sum of the attenuation along a line in the body; that is, anatomic structures are piled on top of each other and flattened into the radiographic image. In an attempt to give a different perspective, one may obtain projections from two different directions (eg, a lateral and an anteroposterior radiograph), so that the structures that are piled on top of each other differ in each projection. In the late 1960s a British engineer, Geoffrey Hounsfield, concluded that if one obtained projection data from a sufficient number of different angles, one could reconstruct the attenuation properties of each volume element in the body and display these as a cross-sectional image. This required the computational power of a computer to accomplish, and the basic idea is illustrated in Figure 2-4. The x-rays from a source are detected by a series of individual detectors (rather than film) after penetrating the body, and each detector defines a ray from the source through the body, thereby creating a projection. The width of the x-ray beam in the dimension perpendicular to the page is only about 10 mm; hence, only one slice of the body in the longitudinal direction is imaged at a time.
Computed tomographic (CT) scanning geometry. A single projection of the head is illustrated.
The x-ray tube and the detector bank are rotated 360 degrees about the patient to obtain, for example, 720 projections at 0.5-degree intervals. The computer is then able to reconstruct a cross-sectional image of the slice of the body by dividing the slice into an imaginary matrix. In a matrix of 512 × 512 pixels in the transverse plane, each pixel represents an area of about 0.5 × 0.5 mm in a 25-cm diameter body. The computer assigns a numerical value to each pixel, which represents the amount of attenuation contributed by the volume element of the body represented by that pixel, and these numbers are converted into a gray-scale image for viewing. In an axial scan series, after one slice is completed, the patient is advanced via a motorized couch by 10 mm in order to image the adjacent slice, and up to 30 slices (images) may be done to reconstruct the anatomy over a 30-cm length of the patient. Newer scanners, called helical (or spiral) CT scanners, use a continuous advance of the patient through the scan beam rather than the stepping couch motion utilized in axial scans, and axial slices are reconstructed by interpolation of data into the slice from a complete rotation. Multislice helical scanners with subsecond rotation times have been developed that collect data for reconstruction of several slices in each rotation; thus, a 30-cm length of patient anatomy can be imaged in 15 seconds or less.