Experimental Study and Optimization of Scan Parameters That Influence Radiation Dose in Temporal Bone High-Resolution Multidetector Row CT

Discussion

Two sources of undesired patient radiation exposure exist in the z-axis direction in MDCT: overbeaming³ and overscanning.⁴ Overbeaming is defined as the penumbra-to-umbra ratio, and it degrades the geometric dose use of a scanner,^3,5 thus exposing a patient to unnecessary radiation; however, wider detector collimations lead to a smaller penumbra-to-umbra ratio and higher geometric dose use.^6,7 In contrast to overbeaming, the effect of overscanning increases with the increasing cone angle of large-coverage MDCT scanners and contributes a larger proportion of the total effective dose for small craniocaudal scan lengths, such as those associated with temporal bone imaging.

In axial imaging, the 2 × 0.625-mm detector collimation had more overbeaming, yielding a geometric dose use of only 37%. Therefore, with the same scanning parameters, the CTDI_vol with the 2 × 0.625-mm detector collimation increased to 47.2 mGy as shown in Table 1, approximately 2 times the CTDI_vol of the other collimations. The 40 × 0.625-mm detector collimation provided the best balance among CTDI_vol, noise, and CNR; however, the 25-mm coverage of the 40 × 0.625-mm detector collimation required 2 rotations to completely image the temporal bone, resulting in an imaged length of 50 mm. In clinical practice, a scan length of 40 mm is typically used to image the temporal bone. The coverage of the 16 × 0.625- and 64 × 0.625-mm detector collimations was 20 and 40 mm, respectively. Thus, the anatomy could be imaged with 1 or 2 rotations, respectively. As a result of the 10 mm of overscan (50-mm scan length versus 40-mm anatomic extent), the DLP associated with the 40 × 0.625-mm detector collimation was 34% higher than the DLP associated with the 16 × 0.625-mm detector collimation and 56% higher than the 64 × 0.625-mm detector collimation. To this end, the 64 × 0.625-mm detector collimation resulted in the lowest overall dose for the examination, though the noise was slightly higher and the CNR slightly lower, respectively, than the 40 × 0.625-mm detector collimation. Nevertheless, it should be noted that the noise and CNR differences between the 64 × 0.625- and 40 × 0.625-mm detector collimations may be less than the differences in the same metrics among subjects in practice, and may reflect the limited sample size in this study. In contrast, one would not expect the dose to vary from subject-to-subject with the same protocol parameters.

Overbeaming and overscanning also occur in helical scanning, though they differ in the relative degree that they influence radiation dose. Overbeaming has a detrimental effect on geometric dose use, particularly at narrower detector collimations such as 2 × 0.5 mm. The higher proportion of overbeaming at this collimation resulted in a substantially higher CTDI_vol than the CTDI_vol from other detector collimations with the same effective tube mAs. In helical MDCT, overscanning caused by the extra data acquisition necessary for reconstruction contributes a larger percentage of the total radiation exposure than overbeaming. For example, the overscan associated with 40 × 0.625- and 64 × 0.625-mm detector collimations was 50.17 and 80.2 mm, respectively. These scan lengths are longer than the 40-mm coverage necessary for temporal bone imaging. As a result, in Table 2, the DLP associated with 64 × 0.625-mm detector collimation was 57.1% higher than the DLP of the 40 × 0.625-mm detector collimation and 78.8% higher than the DLP of the 16 × 0.625-mm detector collimation at a pitch of 1.06. A different choice of pitch would have reduced the degree of overscan, as discussed later in this section. This also highlights the benefit of newer generation scanners that have dynamic z-collimation available to prevent such overscanning.⁸

Another factor influencing the decision of the appropriate temporal bone protocol is the choice of system resolution mode. In helical scanning the default resolution of the 2 × 0.5-mm and 20 × 0.625-mm detector collimations was ultrahigh. Although this resolution mode allowed a substantially higher spatial resolution, it also led to a substantially higher axial and coronal image noise. As a result of this higher noise, we felt that 16 × 0.625- and 64 × 0.625-mm detector collimations were better choices. Given the lower radiation dose and noise, we feel that the 16 × 0.625-mm detector collimation is best for helical temporal bone CT scanning.

Radiation dose is proportional to the square of kVp. For children and smaller adults, lower tube voltages may reduce dose and enable CNR equivalent to that obtained at higher tube voltages.^9–12 In this study, the system had 3 different tube voltage stations: 80, 120, and 140 kVp. Corresponding effective tube mAs values (540, 240, and 180 mAs/section), as shown in Table 3, were available to achieve the same x-ray flux; however, the measured CTDI_vol with 80 kVp was only 14.5 mGy, 57.1% of the CTDI_vol with 120 kVp, and 56.9% of the CTDI_vol with 140 kVp. These differences exist because the adult skull has a stronger ability to attenuate x-rays produced at 80 kVp. In this way, many lower energy (soft) x-rays were absorbed by the skull, and the dosimeter measured a quantum of x-rays that had been attenuated and was substantially less than that of the other tube voltages. As shown in Fig 2, the contrast of the temporal bone structures was such that the bony cochlea and the surrounding tissue in the images produced at 80 kVp was significantly higher than that of the other tube voltages; however, the noise in the reformatted axial and coronal images was substantially higher than with the other tube voltages. The CTDI_vol at 120 and 140 kVp was 25.4 and 25.5 mGy, respectively. When selecting 140 kVp, the noise was lower than the noise at 120 kVp, and the CNR was relatively high. Considering all factors, the best tube voltage for temporal bone CT is 140 kVp.

Tube current has a linear relationship with radiation dose. Tube current should be selected to reduce radiation dose as low as reasonably achievable while still meeting the diagnostic requirements for temporal bone imaging.^13–16 Without such adjustment, particularly for small size or pediatric subjects, patients may receive more radiation dose than is necessary. It is the principal responsibility of CT operators to take patient size into consideration when selecting radiation dose-related parameters, the most commonly adjusted of which is tube current.^17,18 This experiment showed in Table 5 that all 3 readers judged both axial and coronal images to be of diagnostic quality with an effective tube mAs of 120 mAs/section. Increasing the tube mAs above 120 mAs/section raised the image quality rankings, but had no effect on the diagnostic quality of the images.

Pitch is the ratio of the table feed during 1 rotation of the x-ray tube and the beam width in the z-direction, and it reflects the degree of overlap of the radiation beam in helical scanning. In this way, a pitch of 0 indicates that the table has not moved (complete overlap between successive rotations) and a pitch of 1 indicates no overlap between successive rotations. The radiation dose to a patient may decrease as pitch is increased, due to the reduction in the overlap of successive rotations; however, the degree of overscan—and thus the radiation dose to the patient—may be greater with a higher pitch. In our experiment, the DLP was substantially higher with pitches >1. From the aspect of reformatted axial and coronal image quality, noise and CNR were not ideal if the pitch was too small or too large. After comprehensive analysis, it was determined that when the pitch was equal to 0.685, the noise and CNR were at an optimal level and the DLP was minimized.

We acknowledge some limitations of our study. The exsomatized cadaveric heads used in this experiment were stored in a formalin solution for a long period. The muscular tissues and encephalic structures dehydrated noticeably during this time. As a result, there was decreased x-ray attenuation in the cadaveric heads, and the effective tube mAs found here is probably lower than that which would be appropriate in a living human. Our results also do not reflect the further dose reductions that may be enabled with newer dose-saving technologies such as dynamic z-collimation or iterative reconstruction. Also, our results apply only to the assessment of the temporal bone with a particular scanner and would not be directly applicable to eg, other body parts, clinical indications, scanning ranges, structural characteristics, diagnostic requirements, and scanner models; however, the experimental methods and implementation with comprehensive consideration of various dose factors such as detector collimation, kVp, effective mAs, and pitch, may be used as a means to select parameters and to systematically optimize dose for other such examinations.