The following sample essay on “Multidetector CT technology”: revealing the problem of radiation and bad impact of Multidetector CT technology.
The quantities used for radiation dose are absorbed dose which is the energy deposited in tissue/organs per unit mass measured in Gray (Gy) and effective dose which represents a calculated quantity that accounts for the difference in radio-sensitivity of different tissues measured in Sieviets (Sv) (Bae, Hong and Whiting, 2004). The principal dosimetric quantities that are displayed on the MDCT user interface include CT dose index volume (CTDI vol) and dose length product (DLP).
The CTDI integrates the radiation dose delivered both within and beyond the scan volume. CTDI represents average absorbed dose across the field of view for contiguous CT acquisitions and takes into account regional variations in the absorbed dose and it allows for comparison of the radiation doses among different imaging protocols (Wilting, 1999).
When considering image quality it is noted that as image noise decreases there is an increase in radiation dose, however diagnostic information can be attained with significant dose reduction because an increase in radiation dose above a certain level may not contribute to improving diagnostic image quality and merely deposits more radiation into the patients body.
Radiation dose increases roughly with the square of the tube voltage and increases as the tube current and the scan time is increased. These parameters are generally limited in spiral CT protocols, hence the dose is often lower than that in conventional CT (Wilting, 1999). However, tube current and scan time reduction lead to increase in image noise that can adversely affect diagnostic image quality.
So a compromise should always be reached when considering the two.
Radiation dose can be modified by adjusting the tube voltage, tube current, pitch, and scan time. Reduction of the tube voltage at a constant tube current or reduction of the tube current at a constant tube voltage will decrease the output of the x-ray tube and radiation dose to the patient. But an inappropriate reduction of the voltage may result in a marked increase in CT tissue attenuation and noise, particularly in large patients. Instead of manipulating the tube voltage, a more common, practical approach of modifying radiation dose is to adjust the tube current or rotation time. The tube current is usually adjustable in a wide range from 20 500 mA. Another practical advantage of adjusting the tube current instead of tube voltage is that its effect on the image quality is more straightforward. Radiation dose and image noise are affected by the product of tube current and gantry rotation time of scan (Bae, Hong and Whiting, 2004).
Apart from tube voltage, scan time and tube current, CT scanning factors that can be adjusted to optimize radiation dose include gantry rotation time, automatic exposure control, detector configuration, pitch, table speed, slice collimation, scan modes, scan region of interest, scanning phases, post-processing image based filters, metal artefact reduction (MAR) software, and shielding devices. In addition, there are several scan features that users cannot change, including scanner geometry, X-ray beam filters, pre-patient tracking of X-ray tube focal spot, and projection adaptive reconstruction filters (Kalra, n.d.).
MDCT scanners in general have lower geometric efficiency than single-detector row CT scanners because of the gaps between detector elements in the detector array and the use of wider beam widths. The edge of the beam or penumbra has spatially varying x-ray intensity. The penumbra radiation of MDCT scanners does not contribute to image generation but to patient radiation dose. The effect of the penumbra is greatest on four-detector row CT scanners operating with narrow collimation modes. The effect is progressively less with an increase in the number of detector rows, e.g., in 8- and 16-detector row CT scanners, because a fractional effect of the penumbra per detector row becomes proportionally less (Bae, Hong and Whiting, 2004). Pitch and radiation dose are inversely proportional. Scans with a pitch of 2 give 50% of the radiation doses of scans with a pitch of 1 (Zacharias et al., 2013).
Bowtie filters are CT filters that harden the x-ray beam by removing all of the low-energy x-rays that would otherwise be absorbed by the patient and not reach the detector. They also concentrate the x-rays in the central part of the scanned object. This leads to increased image quality and a 50% reduction in surface dose when compared with flat filters. The functionality of bowtie filters depends crucially on the proper positioning of the patient in the gantry isocenter (Zacharias et al., 2013).
Automatic exposure control (AEC), which modulates radiation intensity depending on the patient size, z-axis thickness (Z-DOM) or angular thickness (D-DOM) contributes to dose reduction in MDCT (Lee et al., 2010). Automated Exposure Control (AEC) allows the user to determine the image quality (e.g., noise or contrast-to-noise ratio) requirements, and the imaging system determines the right mAs. In patient size AEC, higher mA is used for larger patients; in Z-axis AEC, higher mA is used at more attenuating positions along the z-axis (patient) and in angular AEC, the degree of modulation depends on the asymmetry of the patient. Another dose reduction method is the in-plane bismuth shield which attenuates radiation to reduce the CT doses of the tissues underneath the shield.
In 4-MDCT systems, a large percentage of the x-ray beam width is wasted when thin (< 2 mm) slices are acquired. This inefficiency becomes small, of the order of few percent, in MDCT with 16 or more detector rows. MDCT systems acquiring 16 or more simultaneous slices should be used, whenever possible, for applications requiring narrow image widths (1 mm or less) to optimize dose efficiency.
When acquiring data in the spiral mode, all CT scanners require an additional rotation or so of data collection at the beginning and end of the scan in order to obtain sufficient data to reconstruct images over the prescribed volume. As the total detector width of MDCT scanners increases or the total scan length decreases, the percentage inefficiency from this effect increases.
As x-ray tube technology has evolved, MDCT scanners have been able to operate at higher power levels, allowing both faster rotation times and longer total scan times. This reduction in the constraints on the x-ray tube in MDCT offers the potential to improve diagnostic image quality, but can also lead to increased doses if care is not taken to optimise scanning protocols.
If the identical mA settings are used for MDCT that were used in SDCT, even for a scanner from the same manufacturer, there can be an unnecessary increase in patient dose. This is primarily due to differences in the distance from the x-ray tube to the patient and detector array, although differences in tube and detector design between the scanner models also play a role. This underscores the fact that the transfer of scanning protocols from one scanner to another should always be performed with caution, so that image quality can be maintained with similar or lower radiation dose depending on scanner characteristics.
When imaging paediatrics, only necessary CT examinations should be performed. When appropriate, other modalities such as ultrasound or magnetic resonance imaging (MRI), which do not use ionizing radiation, should be considered. When it is necessary to use Ct then paediatric CT protocols should be adopted where the beam energy is greatly reduced. The exposure factors are further adjusted based on child size, region to be scanned, the organ system to be scanned and scan resolution. The highest quality images that is those that require the most radiation are not always required to make diagnoses. In many cases, lower-resolution scans are diagnostic. Multiphase examinations result in a considerable increase in dose and are rarely necessary, especially in body (chest and abdomen) imaging when imaging paediatrics and should therefore be avoided.
The most clinical applications of MDCT are in patients with multiple injuries, paediatrics, uncooperative patients, CT colonography, CT angiography, CT bronchoscopy, liver CT, cardiac CT and chest CT. Due to the ability of MDCT to obtain high image quality with minimal or no sedation and at times even despite patient movement it can be used in evaluation of sick patients such as poly-trauma victims, pediatric and uncooperative patients (Aggarwal et al., 2002).
MDCT scanners for coronary imaging utilize a minimum of 16 slice, and now 64 – 320 slice scanners are widely used to get excellent, high-resolution images of the heart and the coronary arteries (Rao and Thompson, 2011). The coronary CT angiogram (CCTA) requires adequate patient preparation and co-operation for good image acquisition. Beta-blockers are routinely given to slow the heart rate in order to eliminate or reduce artefact from motion of the coronary arteries. Temporal resolution is an important limitation and is determined by the speed of the X-ray gantry. MDCT scanners improve this temporal resolution by acquiring a full CT slice in only half a rotation. Images are most commonly acquired using prospective gating and only during diastole (for example at 70-80% of the R-R interval) in order to reduce the radiation dosage. The maximum spatial resolution of MSCT is 0.4 mm, and is determined by the size of the picture elements of the CT detector. Smaller coronary segments frequently are not evaluable with CTA and coronary CTA may be unable to distinguish moderate from severe flow limiting stenosis, because of these limitations of resolution (Rao and Thompson, 2011).
There is a correlation between the amount of calcium present in coronary arteries with the severity of coronary artery disease and thus the likelihood of a cardiac event for example coronary atherosclerosis. Calcium scoring of coronary arteries can be performed with Non-contrast MDCT using either prospective or retrospective electrocardiographic gating. The images thus obtained are then evaluated by software that enables the detection of calcium within the vessels and its volume. A score is then generated that is used to grade the severity of calcification. Since this technique is easily reproducible, it can be repeated to assess temporal regression or progression depending on whether or not the patient is put on appropriate therapy (Aggarwal et al., 2002).
MDCT virtual bronchoscopy is used for non-invasive endoluminal assessment of central airways stenosis. MDCT, unlike single detector CT has the advantage of providing high resolution multiplanar reformatted images because of its high z-axis resolution. Virtual bronchoscopy produces a three-dimensional fly-through view of the tracheobronchial tree from CT data. It takes advantage of the natural contrast between the air-containing lumen and the surrounding tissues. The specific advantage of using virtual bronchoscopy to grade stenosis is that the perspective of the endoluminal view, unlike that in orthoplanar images, stays within the axis of the airway, allowing reliable semiquantitative assessment of tracheobronchial stenosis (Hoppe et al., 2002).
The CT examinations can be performed on a MDCT scanner with collimation, 4? 2mm; pitch, 1.375 (corresponding to manufacturers pitch of 5.5); rotation time, 0.75 sec; 120 kVp; and 100180 mAs. Acquisition time can be roughly 30 sec to allow completion of the acquisition during a single breath-hold. The thorax is scanned during inspiration in a caudo-cranial direction after power injection of 80 mL (flow rate, 2 mL/sec; scan delay, 30 sec) of IV contrast medium containing 300 mg I/mL. The reconstruction intervals and slice thickness are 2 mm.
The basis of CT perfusion imaging is the tracking of a single injected bolus of iodinated contrast material through the cerebral circulation via sequential spiral CT scanning. Using this technique, the following parameters are measured: cerebral blood flow (CBF), cerebral blood volume (CBV), time to peak (TTP), and mean transit time (MTT). CBF is measured in mL of blood per 100 g of parenchyma per minute (normal: 50 mL/100 g/min), whereas CBV is measured in mL of blood per 100 g of parenchyma (normal: 5 mL/100 g) , . MTT is a measurement of the mean time for blood to travel through a given volume of brain, thereby reflecting the amount of time it takes for the bolus of contrast material to pass from the arterial to the venous circulation. TTP is the delay between the first arrival of contrast material intracranially and the time at which the contrast reaches its peak concentration in a given region of parenchyma. Using these parameters, one can utilize hemodynamic differences to assess the intracranial vascular physiology. In cases of normal intracranial perfusion, there is symmetry of all CTP parameters, with CBV and CBF being higher in gray matter than white matter secondary to normal hemodynamic differences between these tissues. In this way, CTP is a form of physiologic imaging, representing active cerebrovascular physiology, rather than just its result e.g., hypo density on a CT scan, reflecting a completed infarct.
One disadvantage of MDCT is a very high image load for the radiologist. A typical abdomen-pelvis study can generate up to 1000 images.
MDCT scanners have progressively increased the number of detectors and reduced scan acquisition times. The latest 64-slice CT systems with gantry rotation times of 0.33 seconds and a spatial resolution of 0.4 mm are now in clinical use. Cardiac imaging, first made feasible with ECG-gated 16-slice CT, will become a routine, highly reproducible examination with 64-slice CT. It is projected that cardiac CT will be the largest growth area for CT imaging over the next five years.
Both 16-slice and 64-slice CT scanners acquire data as isotropic voxels. This means that images can be viewed in all imaging planes with similar spatial resolution leading to routine utilization of 3 dimensional visualization tools. High resolution CT angiography (CTA) for detection of vascular occlusion and stenoses in the carotids, abdominal or thoracic aorta, and the peripheral circulation has replaced conventional diagnostic angiography. CT pulmonary angiography for diagnosing pulmonary embolism has, in most institutions, replaced ventilation-perfusion scans. High resolution MDCT imaging of the liver, pancreas, and kidneys not only improves lesion discrimination but provides a vascular imaging map useful in planning surgical procedures.
As scan times shorten, with 16- and 64-slice CT, it is necessary to delay the onset of scanning in order to image during peak aortic enhancement particularly for CTA applications. Smaller volumes of contrast material may be used but the magnitude of enhancement is reduced, so it becomes necessary to compensate by increasing the injection rate and increasing the concentration of contrast material. Therefore, an ideal technique with a short scan duration decreases the amount of contrast material, increases the injection rate, and uses a high concentration of contrast material.
Additional improvements in arterial enhancement can be achieved by the use of a saline flush. The magnitude of arterial enhancement can be increased with a compact contrast bolus, and contrast volumes may be decreased up to 20%. Contrast material that would remain in the intravenous line or brachiocephalic vein is thereby flushed into the vascular system. Artifacts from dense contrast material in the superior vena cava and right atrium can be eliminated, markedly improving cardiac and pulmonary imaging (Costello, 2006).
Dual source CT or Dual Energy CT (DECT) principle is the acquisition of 2 datasets from the same anatomic location with different kVp (usually 80 and 140 kVp). Using MDCT an entire body can be scanned within seconds using the DECT technique, and thus, misregistration artefacts due to breathing are eliminated as the images are acquired during a single breath hold (Karcaaltincaba and Aktas, 2011).
Neurological applications permit the generation of virtual non-contrast images for the detection of brain haemorrhages in patients who undergo contrast-enhanced CT angiography (CTA). They also allow the removal of bone and calcium from the carotid and brain CTA.
In thoracic applications the major advantages of DECT are a lack of misregistration and visualization of the lung perfusion and ventilation. Misregistration is avoided due to the simultaneous acquisition of 80 and 140 kVp images. In patients with pulmonary thromboembolism, DECT may allow the detection of subtle emboli by revealing perfusion defects.
Clinical applications of cardiac DECT requires 100 and 140-kVp acquisitions. The addition of an iterative reconstruction technique may allow a better image quality for cardiac applications. The acquisition of dual-energy cardiac perfusion combined with coronary CT images can be difficult in patients with elevated heart rates due to a worsening of the temporal resolution from 83 to 165 ms in the dual-source CT. Therefore, dual-energy technique is preferred in patients with low heart rates.