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求助英译汉，急！

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1 X-ray Beam Filtration
The use of an absorbent material between the x-ray tube and the patient can be used to “harden”
the beam such that low-energy x-rays (which contribute disproportionately to absorbed dose) are
reduced, or to “shape” the x-ray beam to deliver dose in the most appropriate spatial distribution.
Previously, only head and body beam shaping (e.g., “bowtie”) filters were available. Recently,
manufacturers have added filters more specific to cardiac imaging or different-sized patients.
2 X-ray Beam Collimation
The use of a very attenuating material between the x-ray tube and the patient should be used to limit
the x-ray beam to the minimal dimensions required. Such collimation occurs along the z-axis to
define the radiation beam width. Additional collimation after the patient to further define the image
width, whether an absorbent material or electronic, causes radiation dose to the patient to be wasted.
Finally, the fan angle of the beam should be collimated to the diameter of the patient to reduce the
amount of bypass that can then be scattered back towards the patient or towards personnel. Such inplane
beam collimation is typically implemented by use of an appropriate scan FOV (shaping filter).
MDCT systems have been observed to have a radiation dose inefficiency at narrow beam collimations,
resulting in a higher CTDI for the narrow beam collimations required for narrow slice
widths. In SDCT, CTDI is generally independent of slice width(although for some SDCT systems,
the CTDI can increase by as much as a factor of 2 for scan widths less than 2 mm).
The dose inefficiency in current MDCT designs is due to unused x-ray beam that strikes outside
of the active area of the detector (along the z-direction). The z extent of this unused portion
of the x-ray beam is approximately constant in size for the various detector configurations; thus
the inefficiency caused by the unused radiation is relatively greater at narrow beam collimations.
In 4-channel MDCT systems, the narrow beam dose inefficiency can be substantial, resulting in
as much as a 40% to 50% dose increase for the narrow beam collimations (4x1 mm or 4x1.25
mm) relative to the widest beam collimations (4x5 mm or 4x8 mm). For submillimeter beam collimations
on 4-channel MDCT systems, this dose increase can be over 100% relative to the widest
beam collimations. The use of a greater number of data channels (16 or more) covering larger zaxis
extents of the detector increases the dose efficiency of MDCT to nearly that of SDCT.
3 X-ray Tube Current (mAs) Modulation and Automatic Exposure Control (AEC)
It is technologically feasible for CT systems to adjust the x-ray tube current (mA) in real-time during
gantry rotation in response to variations in x-ray intensity at the detector, much as fluoroscopic
x-ray systems adjust exposure automatically. This capability, in various implementations, is available
commercially on MDCT systems in response to wide interest from the radiology community. Some
systems adapt the tube current based on changes in attenuation along the z-axis, others adapt to
changes in attenuation as the x-ray tube travels around the patient. The ideal is to combine both
approaches with an algorithm that “chooses” the correct tube current to achieve a predetermined
level of image noise.
By decreasing or increasing the x-ray tube current, the radiation output of the tube is proportionately
changed. Image noise is dominated by the noisiest projection (which corresponds to the most
attenuating paths through the patient). Hence data acquired through body parts having less attenuation
can be acquired with substantially less radiation without negatively affecting the final image
noise. This principle can be applied to modulate the mA angularly about the patient (anteriorposterior
[AP] vs. lateral) as well as along the z-axis (neck vs. shoulders); the tube current can also
be modulated within the cardiac cycle (systole vs. diastole), or with respect to sensitive organs (PA
vs. AP)。
With regard to cardiac CT, the radiation dose for a retrospectively gated exam, where the x-ray
tube is kept continuously on throughout the acquisition, can be dramatically decreased if the tube
current is reduced during portions of the cardiac cycle that are not likely to be of interest for the
reconstructed images. Thus, in addition to modulation of the tube current based on patient attenuation,
the tube current can be modulated by the ECG signal. Since cardiac motion is least during diastole
and greatest during systole, the projection data are least likely to be corrupted by motion artifact
for diastolic-phase reconstructions. Accordingly, the tube current is reduced during systole. Dose
reductions of approximately 50% have been reported using such a strategy. The implementation of
these and other dose reduction strategies is expected industrywide over the next several years, in
response to the strong concern about the radiation dose from CT from both the public at large and
the medical community.
In addition to technical methods of dose reduction, investigators are working to determine clinically
acceptable levels of image noise for a variety of diagnostic tasks. That is, high-contrast exams(e.g., lung, skeletal, colon, sinus) require much less dose (can tolerate higher noise levels) compared
to low-contrast exams (e.g., brain, liver, and other abdominal organs). If the required noise level can
be predefined, CT systems can use technical approaches to deliver the minimum dose required to
achieve the specified noise level. The definition of a robust and standardized noise metric is required,
however, to allow a manufacturer-independent method of prescribing the desired image quality.
4 Size-or Weight-based Technique Charts
Unlike traditional radiographic imaging, a CT image never looks “overexposed” in the sense of
being too dark or too light; the normalized nature of CT data (i.e., CT numbers represent a fixed
amount of attenuation relative to water) ensures that the image always appears properly exposed. As
a consequence, CT users are not technically compelled to decrease the tube-current-time product
(mAs) for small patients, which may result in excess radiation dose for these patients. It is, however,
a fundamental responsibility of the CT operator to take patient size into account when selecting the
parameters that affect radiation dose, the most basic of which is the mAs.
As with radiographic and fluoroscopic imaging, the operator should be provided with appropriate
guidelines for mAs selection as a function of patient size. These are often referred to as technique
charts. While the tube current, exposure time, and tube potential can all be altered to give the appropriate
exposure to the patient, in CT users most commonly standardize the tube potential ( kVp) and
gantry rotation time (s) for a given clinical application. The fastest rotation time should typically be
used to minimize motion burring and artifact, and the lowest kVp consistent with the patient size
should be selected to maximize image contrast.
Although scan parameters can be adapted to patient size to reduce radiation dose, it is important
to remember certain caveats when contemplating such adjustments. First, body regions such as the
head do not vary much in size in the normal population, so modification of scan parameters may not
be applicable here based on head size.
Numerous investigators have shown that the manner in which mA should be adjusted as a function
of patient size should be related to the overall attenuation, or thickness, of the anatomy of interest
as opposed to patient weight, which is correlated to patient girth, but not a perfect surrogate as a
function of anatomic region. The exception is for imaging of the head, where attenuation is relatively
well defined by age, since the primary attenuation comes from the skull and the process of
bone formation in the skull is age dependent.
Clinical evaluations of mA-adjusted images have demonstrated that radiologists do not find the
same noise level acceptable in small patients as in larger patients. Because of the absence of adipose
tissue between organs and tissue planes, and the smaller anatomic dimensions, radiologists
tend to demand lower noise images in children and small adults relative to larger patients. For
body CT imaging, typically a reduction in mA (or mAs) of a factor of 4 to 5 from adult techniques is
acceptable in infants58. For obese patients, an increase of a factor of 2 is appropriate. For head CT
imaging, the mAs reduction from an adult to a newborn of approximately a factor of 2 to 2.5 is appropriate.
Sample technique charts are provided in appendix A. To achieve increased exposure for obese
patients, either the rotation time, or the tube potential, may also need to be increased.
5 Detector Geometric Efficiency
Ideally, all of the photons that pass through the patient should be used in the image formation
process. However, the conversion of photon energy to electrical signal is not a 100% efficient process
(although it exceeds 90% for modern scintillating detectors). New detector materials having even
MEASUREMENT, REPORTING, AND MANAGEMENT OF RADIATION DOSE IN CT

higher absorption and conversion efficiencies are of course desirable, as are detector and signal processing
electronics with very low inherent noise levels. Additionally, the small detector elements
are divided along the detector arc and along the z axis with radiation-absorbing septa (walls).
These septa also provide essential optical isolation between detector elements, but they, along with
the very fine signal transmission wires, create “dead spaces” in the detector and hence waste radiation
dose. As detectors continue to be divided into smaller and smaller discrete elements, the geometric
efficiency of the detector systems must be maintained. One important step in reducing the dead
space has been to attach and route the signal transmission wires underneath each detector element,
instead of between detector elements. However, ongoing reductions in voxel size will likely be limited
by the exponential increase in image noise that would accompany such changes61.
6 Noise Reduction Algorithms
Data processing can be performed on the raw data (in sinogram space) or on already reconstructed
images to reduce image noise. A variety of approaches are possible, all of which seek to smooth out
random pixel variations (noise) while preserving fine detail and structure (signal). With a successful
noise reduction scheme, an image of adequate quality can be acquired with a reduced patient
dose

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• 丁香早知道 - 03.31 湖北 2 名新冠肺炎康复者殴打医院CT技师，警方介入

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抛砖引玉，在线翻译，请高手们修改

1个X射线束过滤

使用吸湿材料之间的X射线管和病人可以用来“硬化”

光束，低能量X射线（其中贡献不成比例的吸收剂量）是

减少，或“形状”的X射线束交付剂量在最适当的空间分布。

以前，只有头部和身体的光束整形（例如，“领结”）过滤器可用。最近，

制造商已经加入过滤器更具体的心脏成像或不同大小的病人。

2个X射线束准直

使用一个衰减材料之间的X射线管和病人应该被用来限制

X射线束的最小尺寸要求。这种准直发生沿z轴为

确定辐射光束宽度。额外的准直后病人进一步定义图像

宽度，是否一个吸水材料或电子，使辐射剂量的病人被浪费。

最后，风扇的光束角应该是平行的直径减少的病人

数额旁路，然后可以分散回到病人或人员。这种平面

光束准直通常是使用适当的扫描视场（整形滤波器）。

多排螺旋CT系统已观测到的辐射剂量低效率在狭窄的光束collimations，

导致较高的CT DI的窄波束collimations需要窄片

宽度。在sdct，CT DI通常是独立的片宽度（虽然一些sdct系统，

在CT DI可以增加多达2倍扫描宽度小于2毫米）。

在目前的多排螺旋CT剂量效率的设计是由于未使用的X射线光束击中外面

地区的积极探测器（沿z方向）。该程度的未使用部分

的X射线束大约是常数大小的各种探测器的配置；因此

效率低下造成的未使用的辐射是相对较大的窄波束collimations。

在4通道多排螺旋CT系统，窄束剂量不足可能是巨大的，从而

多为40%至50%剂量增加的窄波束collimations（4x 1毫米或4x1.25

毫米）相对于宽束collimations（4毫米或非常毫米）。亚光collimations

在4通道多排螺旋CT系统，这一剂量可以增加超过100%相对于宽

梁collimations。使用一个更大的数据通道的数量（16或更多）覆盖较大的扎克思

程度的增加剂量效率的多排螺旋CT接近，sdct。

3个X射线管电流（马来西亚）调制和自动曝光控制（非共体）

这是技术上可行的系统调整X射线管电流（马）在实时期间

龙门旋转响应X射线的强度变化的探测器，多为透视

X射线系统自动调整曝光。这种能力，在各种实现，可用

商业上的多排螺旋CT系统响应从放射社会广泛的兴趣。一些

系统适应管电流变化的基础上衰减沿z轴，其他适应

变化衰减的X射线管绕病人。理想是结合

方法有一个算法，“选择”正确的管电流达到一个预定的

图像噪声水平。

通过减少或增加的X射线管电流，辐射输出管的比例

改变。图像噪声主要是由最热闹的投影（对应到最

衰减路径通过病人）。因此，获得的数据通过本体部分具有不衰减

可以得到大幅度减少辐射不影响最终的图像

噪声。这个原则可以适用于调节马方位对病人（眼

[美]与横向）以及沿Z轴（颈部与肩膀）；管电流也可以

调制的心脏周期（收缩与舒张），或对敏感器官（坝

与美联社）。

关于心脏电脑断层，辐射剂量为回顾性门控考试，在X射线

管持续在整个收购，可以大大降低，如果管

减少电流，在部分的心脏周期，不可能有兴趣的

重建图像。因此，除了管电流调制根据病人的衰减，

管电流调制可以通过心电信号。由于心脏运动至少在舒张

和最大收缩过程中，投影数据是最容易损坏的运动伪影

为diastolic-phase重建。因此，管电流降低收缩。剂量

减少约50%已报告使用这种策略。实施

这些和其他剂量减少战略预计行业在未来几年内，在

针对强有限公司

2011-10-24 16:44