As discussed above, the quality and quantity of X-rays generated within the X-ray tube depend on a number of factors. Increasing the current applied through the filament within the cathode increases the quantity of electrons released by the cathode per unit time (expressed in mA), which in turn increases the number of X-ray photons ultimately released when the electrons are caused to strike the anode. The current might range from 10 mA to 200 mA for pulsed fluoroscopy and from about 100 mA to 1,000 mA for cine angiography . The kV represents the voltage difference between the anode and cathode. The higher the kV, the faster the electrons are accelerated toward the anode, therefore releasing X-ray photons of higher energy upon collision with the anode. The typical setting of X-ray tube voltage for cardiac fluoroscopy is about 70 kV . Another characteristic of the X-ray tube that affects the X-ray beam is the pulse duration. In pulsed fluoroscopy (as well as in cine) the X-ray beam is pulsed on and off rapidly (e.g., at 30, 15 or 7.5 pulses per second), which results in lower total radiation dose than conventional continuous fluoroscopy. In pulsed fluoroscopy the time interval between pulses does not generate an image, and in order to reduce the impression of flicker of the image, the monitor displays the previously stored image until the subsequent pulse of X-rays generates a new image. Lowering the pulse frequency rate lowers the radiation dose proportionally. The duration of each pulse is also known as the exposure time, and is expressed in milliseconds (ms), with typical settings for pediatric cardiac fluoroscopy ranging from 1 ms to 4 ms per pulse .
Alterations in any of the characteristics of the X-ray beam (mA, kV, and ms) will result in specific changes to the image produced and also result in changes to the radiation dose. For example, increasing the mA will result in more X-ray photons (of the same energy range) in the beam, which in turn produces more darkening of the film. However, an increase in the mA results in a proportional increase in radiation dose. In order to illustrate the interaction of the various beam characteristics with the AEC, let us consider the following example. During a procedure, if the frontal II is changed from a straight posteroanterior projection to a sharply angulated one (e.g., a cranially angulated left anterior oblique projection), the thickness that the X-ray beam will need to traverse will be greater. It is estimated that increasing the tissue thickness by 3– cm results in an image that is half as dense (a 3-cm tissue thickness is considered one ‘half-value layer–. The AEC will, therefore, need to compensate in some fashion to correct the image density. Although doubling the pulse width (exposure time) would effectively result in a correction of the image density, this would also double the radiation dose per frame produced. Furthermore, longer pulse widths might result in images that exhibit motion artifact, rendering them less sharp overall. Doubling the mA would be another way to compensate for the increased tissue thickness, but this would also result in a doubling of the radiation dose, though it would also improve image contrast. The final option would be to increase the voltage by 6–0 kV. This would result in the production of higher energy X-ray photons, which would result in improved tissue penetration. Because fewer photons would be absorbed by the patient, this would actually result in a lower radiation dose to the patient compared to increasing the mA alone, at the cost of decreased image contrast. The purpose of the AEC system is to automatically adjust these parameters in order to maintain a relatively constant image density. AEC algorithms are different among different manufacturers, such that older units might only vary the mA, while newer ones vary both the mA and kV to compensate for a change in image brightness.
Filtration of the X-ray beam is an important way to reduce the patient’s radiation dose. Low-energy X-rays do not contribute to image formation and are instead absorbed in the patient’s superficial tissues. For this reason, filtration of these photons (also known as ‘hardening–the beam) using a layer of aluminum (and in some units, copper is also used) at the exit port from the X-ray tube is beneficial, with minimal effect on image quality.
The X-ray beam is further altered by the use of collimators and partial-thickness filters. Collimators composed of lead shields are used to shape the beam at the exit port from the X-ray tube. The tighter the collimation, the less divergent the X-ray photons will be, which in turn reduces the amount of scatter. Reduction of scatter is extremely important because scattered radiation is the main cause of radiation exposure to the patient’s body outside the field-of-view and to the personnel in the room. Scatter is also an important cause of image degradation; therefore, tight collimation on the region of interest also serves to improve image quality. Another benefit of tight collimation is that less volume of the patient’s body is exposed to the primary beam (although reduction of scatter exposure to the patient’s tissues is important, the primary beam is many times more intense than scattered rays). A qualification with respect to the benefits of tight collimation is in order, however: if collimation is extremely tight, part of the AEC sensor (typically located at the center of the image detector) might lie within the collimated region and will thus perceive reduced brightness. In such situations AEC compensation will result in an image that is excessively bright. Reducing the collimation somewhat will correct the problem.
Having discussed the characteristics of the beam at the X-ray tube, let us now turn our attention to the II. An important (and under-utilized) radiation reduction strategy in pediatric fluoroscopic imaging is the removal of the antiscatter grids. The grids are composed of lead strips designed to filter out scattered X-ray photons. Because the amount of scatter is greater with larger patients, the grids are most beneficial when used during studies on large patients. Conversely, their use in studies on young children results in relatively little improvement in image quality and, in fact, results in increased patient doses. This is because the grids also filter out some of the non-scattered rays, resulting in decreased image brightness; AEC compensation, therefore, leads to increased radiation dose in order to maintain image brightness at an acceptable level. Therefore, when used in studies on young children, the grids represent a cause of increased radiation exposure without significant benefit in image quality. Removal of the grids should be considered in these cases. One potential problem is that not all catheterization laboratories are equipped with grids that are easily removable. Even among our own laboratories at Texas Children’s Hospital there is variability in the design of the grid attachments such that one unit has grids that can be removed in about 5 minutes with minimal effort, while another requires significant dismantling of the II unit and requires approximately 30 minutes to accomplish. Because pediatric catheterization laboratories care for patients with an extremely wide range of weights and sizes, the easy removal and replacement of antiscatter grids in between cases might not always be possible. However, a case that is expected to require a long fluoroscopy time in a small child should warrant the consideration of removal of the grid when this is feasible.
The age of the II also affects radiation dose because the efficiency of conversion of X-ray photons into light photons at the cesium iodide input phosphor decreases with time. The impaired efficiency with advancing age causes a compensatory increase in radiation dose. Furthermore, the image quality also degrades with advancing age. Competent biomedical engineers working with a radiation physicist must regularly evaluate catheterization laboratories for image quality and dose requirement.
An understanding of the various image magnification techniques is necessary to achieve optimal diagnostic accuracy while reducing radiation dose. The least costly method (in terms of radiation dose to the patient) of magnifying an image is the use of the ‘replay zoom–feature. This feature allows an image acquired in a non-magnified fashion to be displayed on the monitors in a magnified view. Using the replay zoom feature produces an image that does not contain more information (pixels), but rather is simply displayed by replacing each pixel with a larger block on the screen. The end result is an image that is noticeably coarser and less pleasant to view. Reviewing an image in replay zoom mode can be helpful when making measurements of a small vessel, for example, because positioning the digital calipers on the edges of the vessel on the screen using the joystick is made easier when the target is amplified. However, vessels too small to accurately see in the standard view will not be viewed more easily in the replay zoom mode. A second type of magnification is called geometric magnification. This is accomplished by moving the II farther from the patient. As the II is moved farther from the patient, the shadows of the vascular structures cast upon it will be larger (analogous to moving a screen farther away from a television projector). The disadvantages of geometric magnification are significant. First, the resulting image will be less bright and will thus prompt a compensatory increase in radiation dose by the AEC system. Second, the image will be affected by increased geometric blur, such that the edges of the vessels will be significantly less sharp. Last, the II, itself, actually helps to block some of the scattered rays to the personnel and to the operator, and moving the II away from the patient removes this protective effect. Overall, the use of geometric magnification is discouraged for all of these reasons.
The last type of magnification is called electronic magnification. This feature is particularly useful in pediatric patients because of the small anatomic structures that are being imaged. Electronic magnification is produced by reducing the field-of-view, such that a smaller surface area on the II is being exposed, with the resulting smaller exposed area being magnified to fill the entire visible screen. Most IIs have two or three ‘modes– or fields-of-view. For instance, a 9-inch II might have a 9-inch, a 7-inch, and a 5-inch mode (roughly equivalent to 23 cm, 17 cm, and 13 cm). Because a smaller area of the II receives the transmitted radiation, a loss of image brightness occurs, resulting in a compensatory increase in radiation dose. The increase in radiation is substantial, and thus electronic magnification should be used sparingly and only when truly necessary. The maximal increase in entrance exposure to the II is calculated to be the ratio of the area of the standard view over the area of the magnified view  (e.g., in going from a 9-inch mode to a 7-inch mode, the exposure is increased by a factor of 92/72=1.7, while going from a 9-inch mode to a 5-inch mode increases the exposure by a factor of 92/52=3.2). This holds true if only the mA is increased when going to a magnified mode. With changes in kV also possible on some units, the actual increase in dose is sometimes significantly less than described by the above formula. Nonetheless, an increased dose is certain. However, the increased radiation cost to the patient is accompanied by an increase in image sharpness and resolution, thereby improving the visualization of small vascular structures. It is important to remember that an entire procedure need not be performed in a magnified mode: a critical angiogram can be acquired in magnified mode and the remainder of the procedure can then be carried out in the standard mode. Another disadvantage of overuse of the magnified mode is that excessive panning might be required to find adjacent structures that are out of the field-of-view. This is (unfortunately) commonly the case during coronary angiography, where the table might need to be moved (panned) during the angiogram to see the entire coronary tree. In the pediatric laboratory, where angiograms are almost always performed in biplane, and often in angulated views, resorting to panning to find the structures of interest during the course of an angiogram is almost never satisfactory because movement of the table in order to capture the desired region in one plane often results in loss of the region of interest in the other plane. Instead, a lesser degree of magnification should be used in order to avoid the need for panning altogether. As a general rule, the lowest acceptable level of magnification should be used that still allows sufficient diagnostic accuracy.