Use of Ionizing Radiation in Screening Examinations for Coronary Artery Calcium and Cancers of the Lung, Colon, and Breast Shuai Leng, PhD, Carrie B. Hruska, PhD, and Cynthia H. McCollough, PhD

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ost medical imaging tests are used for diagnostic purposes to evaluate symptomatic patients, but a select few are commonly used for screening, including computed tomography (CT) for coronary artery calcium screening, lung cancer screening, and colon cancer screening; and mammography for breast cancer screening. All of these tests use ionizing radiation. Concerns about potential risk associated with ionizing radiation have increased in recent years, with the increased use of medical imaging. Given that screening examinations are performed on asymptomatic patients, controlling radiation dose to as low as reasonably achievable in these examinations is especially important. The purpose of this article is to discuss issues associated with ionizing radiation in common imaging screening examinations, including metrics used to describe radiation dose, factors for variability in dose among patients, typical dose values in each type of screening examination, dose-reduction methods, and a discussion of risks associated with radiation from these tests. In a previous article by McNitt-Gray in 2003, radiation dose issues in CT screening examinations have been reviewed.1 This article expands to include other screening examinations, such as breast cancer screening. Newly introduced dose metrics and dose-reduction techniques in the past 10 years are also included in this article.

Dose Metrics Multiple dose metrics have been used to quantify ionizing radiation, some of which are generic dose metrics that apply to all imaging modalities, whereas others are more modality specific. In this section, the definition of these dose metrics is reviewed so that they can be appropriately used in the right context. Methods to measure or estimate these dose metrics from each imaging modality are also reviewed. Department of Radiology, Mayo Clinic, Rochester, MN. Address reprint requests to Shuai Leng, PhD, Department of Radiology, Mayo Clinic, 200 First St SW, Rochester, MN 55905. E-mail: leng.shuai@mayo. edu

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http://dx.doi.org/10.1053/j.ro.2014.10.012 0037-198X/& 2014 Elsevier Inc. All rights reserved.

General Dose Metrics A list of common dose metrics and their use in radiation dosimetry is given in Table 1. Exposure is a common metric used to quantify the ability of x-rays and gamma-rays to ionize air or create charges. Exposure is expressed as charges per unit air with the SI (International System) unit of Coulombs per kilogram (C/kg), or the non-SI but still commonly used unit, Roentgen (R), where 1 R ¼ 2.58  104 C/kg. Exposure can be measured using an ionization chamber in air, in phantoms or on the body surface. An alternative dose metric is air kerma. Kerma stands for “kinetic energy released in matter” and refers to the radiation energy that is transferred to charged particles per unit mass of material, with SI units of joule per kilogram (J/kg) or Gy, where 1 Gy ¼ 1 J/kg. For medical imaging, Gy is a rather large quantity and milligray (mGy) (103 Gy) is a more common unit. When kerma is measured in air, it is referred to as air kerma and indicates energy absorbed in a kilogram of air, as radiative energy loss is negligible in air. A measured exposure of 1 R in air gives air kerma of approximately 8.7 mGy. Air kerma is used in most countries outside the United States. Absorbed dose is defined as the total energy deposited in unit mass at a small local area and also has units of J/kg or Gy. Absorbed dose in individual organs (organ dose) is very useful to quantify radiation dose associated with the imaging examination and the potential dose risk associated with the radiation. Direct measurement of absorbed dose can be performed on size-matched anthropomorphic phantoms, with structures simulating human organs, to provide an estimate of patient organ dose with the same scanning techniques. To estimate organ dose from patient examinations, Monte Carlo simulations are usually used, which can simulate photon transmission and dose deposition based on x-ray output and patient attenuation properties. A limitation of Monte Carlo simulation is that it is usually time and labor intensive. Therefore it is mainly used in research, not routinely used in clinical practice for individual patients. A limitation of absorbed dose is that it does not take into account the biological effect of the type of radiation. Given the

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Table 1 A Summary of Basic Metrics Used in Radiation Dosimetry Metric

SI Units

Exposure

C/kg

Non-SI Units

What it Represents

How it Is Determined

Roentgen The quantity of ionization (charges) produced Measured with an ionization chamber at point (R) per unit mass (kg) of air when x-rays and of interest. gamma-rays travel through air. The effect of radiation in tissue will be proportional to this ionization in air. Air kerma J/kg or Gy rad Kerma ¼ “kinetic energy released in matter” Measured with ionization chamber calibrated to display air kerma. Air kerma refers to the energy absorbed in a Or measured exposure is converted to air unit mass (kg) of air when ionizing radiation kerma: A measured exposure of 1 R will give travels through air. air kerma of approximately 8.7 mGy at that point. Absorbed J/kg or Gy rad The energy imparted by radiation per unit mass Conversion factors determined using Monte dose (kg). Carlo simulations allow calculation of absorbed dose to organs based on exposure or air kerma. Weighting factors: Equivalent Sv rem The absorbed dose, with adjustment by a dose weighting factor to reflect the type of radiation. Equivalent dose ¼ absorbed dose x weighting x-Rays and gamma-rays ¼ 1 factor Protons ¼ 2 Neutrons ¼ 5-20 Alpha particles ¼ 20 The equivalent doses for each organ exposed, Effective Sv rem The combined detriment from stochastic with adjustment by a weighting factor to dose effects due to the equivalent doses in all reflect the radiosensitivity of each organ, are organs and tissues of the body. Includes summed. Weighting factors are given in potential noncancer detriment. ICRP Report 103. ICRP, International Commission on Radiological Protection. 1 R ¼ 2.58  104 C/kg. 1 Gy ¼ 1 J/kg. 1 rad ¼ 10 mGy. 1 rem ¼ 10 mSv.

same absorbed dose, biological detriment is much higher for protons, neutrons, and alpha particles compared with photons (x-rays and gamma-rays). To take this into account, equivalent dose is used, which is the product of absorbed dose and a radiation weighting factor. The SI unit of equivalent dose is Sievert (Sv), which is also equal to J/kg, but differs from Gy because of the weighting factor applied. For x-rays and gamma-rays, the weighting factor is 1, so for all diagnostic imaging, the equivalent dose has the same value as absorbed dose, although the unit is different. A second limitation of absorbed dose is that it does not account for varying radiation sensitivity among biological tissues. For example, the lung and the breast tissues are more sensitive to radiation than the brain tissues. Tissue weighting factors accounting for these differences have been assigned by the International Commission on Radiological Protection.2 These weighting factors represent the radiation detriment averaged across an entire population, for all ages and both the genders. The sum of the weighting factors for all organs and tissues equals 1. By multiplying the equivalent dose of each organ or tissue by its tissue weighting factor and summing for all organs, a metric called effective dose is determined. The unit of effective dose is Sv, the same as that of equivalent dose. Effective dose can be used to represent the equivalent whole-

body radiation from examinations that involve only partial irradiation and can therefore be used to compare radiation dose between imaging modalities that expose different areas of the body, such as CT and mammography. It is important to note that effective dose applies to a general population, not to a specific individual. Effective dose should never be used for the purpose of assigning risk to an individual.2

CT-Specific Dose Metrics CT dose index (CTDI) and its derivatives, such as weighted CTDI (CTDIw), volume CTDI (CTDIvol), and dose length product (DLP), have been used for many years to quantify CT scanner radiation output in quality assurance programs. The measurement of CTDI is well described in the literature and has been standardized through national regulatory and international standard organizations.3-5 All modern CT scanners are capable of reporting CTDIvol and DLP, which are reported for each examination and can be archived. CTDI is measured using a 10-cm pencil chamber in axial scan mode, with CTDI100 calculated as Z 50 mm 1 DðzÞ dz CTDI100 ¼ NT 50 mm

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150 where N is the number of detector channels along the z-axis, T is the width of each channel, and D(z) is the radiation dose profile along the z-axis. CTDI is measured on a cylindrical phantom made of polymethylmethacrylate that is 14- or 15cm long, and either 16 cm or 32 cm in diameter. Measurements are made at both the edge and the center of the phantom. A weighted average of the CTDI100 at the center (one-third) and the edge (two-thirds) is defined as the weighted CTDI, or CTDIw. To take into account the influence of helical pitch in helical scanning or the spacing in axial scanning, volume CTDI (CTDIvol) was introduced ( CTDIw for helical scan pitch CTDIvol ¼ CTDIw  NT for axial scan I where I is the table translation in 1 rotation. N and T are the same as defined earlier. The unit for CTDI100, CTDIw and CTDIvol is mGy. CTDIvol represents the multiple-scan average dose, but it does not indicate total energy imparted, as scan length is not taken into account. DLP, from which energy imparted can be calculated, is the product of CTDIvol and scan length, DLP ¼ CTDIvol  Scan Length and the unit is mGy · cm. An important caveat of CTDI is that it is not equivalent to patient dose.6 CTDI is measured on standard phantoms, as described earlier, without accounting for patient size and attenuation variation. For the same scanner output (ie, CTDIvol), larger patients have higher attenuation and larger mass, and consequently receive less dose per unit mass. A new concept, named size-specific dose estimate (SSDE), takes into account the influence of patient size and attenuation on CT dose.7 SSDE is used to estimate the average dose in the central region of the scan range based on the product of CTDIvol and a conversion factor based on the patient size. This conversion factor was determined using the Monte Carlo simulations and physical phantom measurements by several research groups. It was found that the conversion factor decreases with increased patient size following an exponential function. Studies have demonstrated that SSDE accurately represents the average dose in the center portion of the scan region (maximum differences from other dosimetry methods are less than 20%). SSDE is not an organ-specific dose, as it represents the average dose in the central region of the scan, not in a specific organ. For large organs located centrally in the scan, it provides a reasonable estimate of organ dose (eg, liver, lung, and colon). However, doses to smaller organs, organs at the periphery of the scan regions, or outside the scan regions are also of interest. Absorbed dose to any specific organ or organs can be estimated by the Monte Carlo simulations using patient CT images, from which effective dose can then also be estimated. Another frequently used method to estimate effective dose in CT is to use the so-called k factor method, E ¼ k  DLP where k is a factor (unit, mSv · mGy1 · cm1) whose values depend on the body region that is imaged.8 However, these estimates of effective dose do not provide risk information for a

given patient, even if the patient-specific organ doses or DLP values are used; the organ and tissue weighting factors and the k-factors are all derived for a population and represent an average over all ages and both the genders. As such, effective dose is never a patient-specific dose or risk metric.

Mammography-Specific Dose Metrics Monitoring of radiation dose is an important and required component of mammographic system quality control. In the United States, mammography is the only medical imaging test with a regulatory limit on radiation dose, per the 1992 Mammography Quality Standards Act (MQSA). Mammography exposes only the breast to ionizing radiation, with negligible scattered radiation received by adjacent organs. Described in simple terms, the breast comprises 3 types of tissue: skin, adipose (fat), and glandular tissue, which absorb radiation dose during a mammogram. Because glandular tissue is the most attenuating component and is thought to be the component at highest risk of radiation-induced carcinogenic effects, an estimate of the average amount of radiation absorbed in glandular tissue was chosen as the metric for mammography dose. This metric is referred to as mean glandular dose or average glandular dose (AGD) and is given in units of mGy. As part of routine quality control for mammographic systems, AGD is measured with a standard phantom meant to simulate an average breast. A variety of breast dosimetry protocols exist worldwide using a variety of phantom designs.9 AGD is not measured directly, but is calculated from a measure of exposure (or air kerma) obtained at the entrance surface of the phantom. A number of Monte Carlo models have been developed to simulate dosimetry of various compressed breast thicknesses and compositions, accounting for the x-ray beam characteristics (tube peak kilovoltage [kVp] and half-value layer [HVL]), for all currently available anode and filter combinations.10-12 From these simulations, tables of conversion factors, called the normalized glandular dose coefficients, were determined for converting measured entrance air kerma (or sometimes entrance skin exposure) to AGD, given by AGD ¼ KE  DgN where KE is the entrance air kerma, and DgN is the normalized glandular dose coefficient for a particular acquisition setup. The breast dosimetry quality control protocol commonly utilized in the United States uses the American College of Radiology (ACR) mammography accreditation phantom, which is a 4.2-cm-thick block with an insert simulating tissue composition of 50% adipose and 50% glandular tissue. Each mammography system manufacturer provides a standard procedure for consistent measurement of entrance air kerma at the entrance surface of this phantom and provides appropriate look-up tables for conversion factors to AGD. MQSA dictates that for the clinically used technique, the AGD to a phantom simulating 50-50 glandular-adipose composition must be less than 3 mGy (300 mrad); mammography cannot be legally performed in the United States on a system that exceeds this limit.

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Factors Affecting Patient Dose Radiation dose received by patients during imaging examinations (screening or diagnostic) depends on both the imaging techniques and the patients’ size and attenuation properties. Common factors affecting patient dose in radiographic and CT examinations include tube potential (kV), tube current (mA), exposure time, and beam collimation. In general, patient dose increases quadratically with tube potential and linearly with tube current and exposure time. For CT examinations, other scanning parameters such as rotation time and helical pitch also affect patient dose. Reconstruction parameters, such as reconstruction algorithms, kernels (or filters), and slice thickness, although do not directly affect radiation dose, have significant effect on image quality. Consequently, these parameters affect patient dose in an indirect way. Patient size and attenuation have direct effect on image quality and radiation dose of screening examinations. To maintain the same image quality, scanner output is usually increased for large size patients to accommodate the increased attenuation so that enough photons can reach the detector. For example, a large adult patient can be scanned with a technique of CTDIvol 10 times higher than the one used for a small pediatric patient. Similarly, a larger size breast with higher density requires higher output from mammographic unit to maintain the same image quality as a regular size breast. However, the increased scanner output does not directly translate to increased patient dose. Dose represents the total energy deposited per mass of tissue. Although total imparted energy increases with the increased scanner output for large size patients, the mass of tissue also increases. Therefore, it cannot be directly concluded that larger size patients are imaged with higher dose. A recent study by Christner et al evaluated 545 adult CT scans of the torso that were acquired with automatic exposure control to adapt the scanner output to patient size. The study demonstrated that although scanner output (quantified by CTDIvol) increased linearly with patient size, patient dose (quantified by SSDE) was independent of size.13

Dose Variation As mentioned in the previous section, patient size and attenuation play a critical rule in image quality and radiation dose. Therefore, for the same scanner, imaging techniques should be adjusted accordingly to obtain the same image quality for all patients. This may result in variations in scanning parameters and scanner output (ie, CTDIvol) that are necessary to achieve comparable diagnostic performance. Differences in scanning techniques unrelated to patient size do exist on the same scanner model, patient population, and imaging task, due to either site preferences or nonoptimized scanning protocols. These variations should be restricted to be as small as possible.

Radiation Dose in Coronary Artery Calcification Screening Coronary artery calcification (CAC) quantified from CT images has been used as a surrogate of atherosclerosis. An example

Figure 1 Example of coronary artery calcification screening examination acquired on a 64-slice dual source CT scanner, with calcium present on the coronary artery (arrow).

image of CT CAC is shown in Figure 1. A special challenge of cardiac CT is the motion of the heart and coronary arteries, which requires high temporal resolution and electrocardiography (ECG)–gating techniques. In addition to the usual factors that affect patient dose, gating techniques, ECG pulsing (will be discussed in the next section), temporal resolution requirements, and patient heart rate can affect radiation dose in CAC. Kim et al14 reviewed the CT CAC literature and estimated the effective dose using a Monte Carlo–based dosimetry and the reported protocols. It was found that dose from a single CAC CT scan varied more than 10-fold (effective dose range ¼ 0.8-10.5 mSv), depending on the scanner and protocol.14 The 2 highest effective doses (8.0 and 10.5 mSv) were from retrospectively gated scans performed without ECG pulsing. All the rest were less than 3.6 mSv. The American Heart Association recommends 1 mSv for prospectively triggered CAC scans and 3 mSv for retrospectively gated scans.15 The Society of Cardiovascular Computed Tomography and the Society for Atherosclerosis Imaging and Prevention Tomographic Imaging and Prevention Councils also issued guidelines on radiation dose and dose-optimization strategies.16,17 New scanning and reconstruction techniques have been developed after the release of these reports and guidelines, which result in lower radiation dose levels. These are reviewed in the section on dose reduction.

Radiation Dose in Lung Cancer Screening Multiple studies have been performed for lung cancer screening using either radiography or low-dose CT. The most recent large-scale study on lung cancer screening is the

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152 National Lung Screening Trial (NLST), in which more than 53,000 persons at high risk for lung cancer were randomly assigned to 3 annual screenings with either a single-view chest radiograph or a low-dose CT.18 Results demonstrated that low-dose CT screening reduced overall mortality by 20% compared with radiography.19 There were 2 physicist quality assurance groups overseeing the study, and they established a common set of protocols in consultation with NLST radiologists. Parameter ranges that were considered suitable for producing acceptable images from the radiographs and low-dose CT examinations were determined and distributed to enrollment sites.20 Chest radiographs were performed on film screen, computed radiography, and digital radiography systems, depending on the availability at the sites. Kruger et al21 estimated the dose in radiographic examinations in NLST. He estimated that the effective dose of a single-view posteroanterior chest radiographic examination was 0.052 mSv, which was within the range reported in the literature (0.30-0.016 mSv).21 Compared with chest radiography, CT has the advantage of generating three-dimensional (3D) tomographic images without anatomical overlay, as occurs in the 2D projection radiographic images. An example image from a low-dose CT performed for lung cancer screening is shown in Figure 2. In the NLST study, half of the patients were scanned on CT scanners with 4- to 64-detector row. The main goal of lung cancer screening is to detect lung nodules, which have high contrast with the air background. This high contrast enables CT scans using lower radiation dose. CT dose was evaluated for a single low-dose CT screening examination performed in NLST based on a “standard man” model and a range of reported techniques used for average-size participants.20 The average CTDIvol was 2.9 mGy, with a standard deviation of 1.0 mGy and a range of 1-8 mGy. These data were averaged by the scanners, not weighted by the number of participants that used a particular scanner. The average effective dose was

1.4 mSv, with a standard deviation of 0.5 mSv, and a range of 0.5-4 mSv. This dose level was substantially lower than standard chest CT examination, whose effective dose was estimated to be 7 mSv (range: 4-18 mSv).22

Radiation Dose in CT Colonography Screening Because of the high tissue contrast between colon and insufflated gas, substantial dose reduction can be obtained without sacrificing polyp detection, which is the main goal of colon cancer screening. An example image of CT colonography (CTC) screening is shown in Figure 3. Multiple screening trials using low-dose CT techniques, such as the Department of Defense trial and National CT Colonography trial, have demonstrated the effectiveness of CTC compared with optical colonography.23,24 In a survey conducted by Liedenbaum et al,25 it was found that the median effective dose was 5.7 mSv (2.8 mSv supine and 2.5 mSv prone) for the CTC screening protocols. This was significantly lower than the daily practice CTC protocols, which had a median effective dose of 9.1 mSv (5.2 mSv supine and 3.0 mSv prone). It was also found that doses did not differ significantly between CT scanners with different numbers of detector rows. Based on these studies, the ACR updated its 2005 guidelines in October 2009 for the performance of CTC. Guidelines on CT scanning techniques have been included.26,27 Patients are scanned using a thin-section, low-dose technique in both the supine and the prone positions on scanners equipped with Z 4-row detectors. Total CTDIvol for both the supine and the prone scan series should not exceed half of the diagnostic reference level for routine abdominal pelvic CT (25 mGy). That is, a CTC examination of a standard-sized adult should have a CTDIvol of no more than 12.5 mGy for both the positions, or 6.25 mGy per position.26,27

Radiation Dose in Breast Cancer Screening Mammography Dose

Figure 2 Low-dose CT lung cancer screening shows a lung nodule in the right upper lobe (arrow).

The United States Preventive Services Task Force currently recommends biennial screening mammography in women aged 50-74 years28; however, most major medical societies, including the American Cancer Society, disagree with this recommendation and continue to recommend annual screening mammography in all women 40 years of age and older. Unfortunately, the benefits vs risks of mammography— detecting cancer before metastasis and reducing breast cancer mortality vs missed cancers, false-positive findings, and overdiagnosis of indolent tumors—are under continual debate,29,30 causing patients and providers to question at which age screening mammography should start, how often to screen, or if it should be performed at all. The low hypothetical risk associated with radiation doses of current mammography equipment, however, should not be a reason to avoid mammography.31 Sample screening mammograms performed in 2 different patients are shown in Figure 4.

Ionizing radiation in screening examinations

Figure 3 CT colonography shows a flat (cigar-shaped) mass in ascending colon, just superior to the level of the ileocecal valve, measuring 1.5 cm greatest dimension (arrow). (Color version of figure is available online.)

Figure 4 Screening mammograms performed in 2 different patients. Mediolateral oblique projections are shown. (A) Invasive ductal carcinoma (arrow) detected in a patient with almost entirely fatty breast composition. (B) Invasive ductal carcinoma (solid arrow) detected in a patient with heterogeneously dense breasts, with composition of approximately 30% glandular tissue. A benign area of calcification was also detected (dotted arrow). (Color version of figure is available online.)

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154 Although doses reported from the early days of mammography in the 1960s and 1970s were relatively high (average AGD of 12 mGy, but as high as 90 mGy) and varied substantially owing to a wide range of techniques in use, modern day mammography equipment delivers very low radiation doses to patients (average AGD of 2 mGy, range of 1-8 mGy) that are considered safe for routine screening in healthy patients.32 Dose to a standard mammographic phantom, as described earlier, with its mandated limit of 3-mGy AGD, does not represent dose to an individual patient. Several surveys have shown wide variability in both compressed breast thickness and breast composition among patients that consequently leads to variability in patient dose.33,34 In a study of more than 2800 women, Yaffe et al reported average compressed breast thickness of 5.9 cm (standard deviation ¼ 1.6 cm) and showed that the 50-50 glandular-adipose composition used in standard phantoms is not representative of that found in patients. Using an algorithm based on volumetric imaging from dedicated breast CT to assess breast composition in mammograms of more than 2800 women, average percent glandular tissue was 19%, with a range of 14%-26%.33 In work conducted to assess automated tools for measuring breast density on mammograms, the average proportion of fibroglandular tissue was 19% and 25% with 2 different algorithms.35 In practice, patient doses from mammography are not routinely performed or reported, but can be estimated by accounting for factors such as the patient’s actual compressed breast thickness, an estimate of the patient’s breast composition, the specific technique used (tube peak kilovoltage and mAs), and measurements of entrance exposure and beam half-value layer obtained at these settings.36 In an analysis by Kruger and Schueler of nearly 25,000 screen film mammograms in more than 6000 women, a set of phantoms with varying thickness and glandular composition were used to simulate the range of patient characteristics observed. By replicating the exposure factors for each mammogram performed, glandular composition for each patient could be estimated, and with application of appropriate conversion factors, patient-specific mean glandular doses were calculated; the AGD per mammographic view was on an average 2.6 mGy (standard deviation ¼ 1.1 mGy).34 Other patient-specific dose surveys for screen film mammography have reported average AGDs of between 1.2 mGy and 2.8 mGy per view.10,34 Reported doses from full-field digital mammography (FFDM) are generally lower than that obtained from screen film systems, primarily owing to the ability to use a harder x-ray beam with a digital receptor compared to what can be used with a screen film cassette.37 A harder beam refers to a beam with higher average beam energy, which is more penetrating. Conversely, a softer beam contains more lowenergy x-rays that are absorbed in breast tissue, and thus contribute to dose but not to image formation. Doses from FFDM and screen film mammography performed in the same patient, as part of the American College of Radiology Imaging Network digital mammographic imaging screening

S. Leng et al trial, were compared in more than 4300 women. On average, the calculated AGD per view was 2.4 mGy for screen film vs 1.9 mGy for FFDM examinations. The dose to each patient, accounting for 2 views (craniocaudal and mediolateral oblique) and any additional views that were acquired, was on an average 4.98 mGy for screen film and 4.15 mGy for FFDM.37 Using the weighting factor for breast radiosensitivity of 0.12,2 these AGDs correspond to effective doses of approximately 0.6 mSv for screen film and 0.5 mSv for FFDM. Effective dose should not be used for cancer risk estimation, particularly for mammography, where only breast tissue is irradiated. Individual risk estimates, if desired, should use the Biological Effects of Ionizing Radiation VII (BEIR) methodology. Efforts to more accurately estimate patient-specific dose for mammography are ongoing, in anticipation of new requirements and demand for tracking of doses in medical records. An automated software program for estimating mammographic breast density has recently been implemented by incorporating individual glandular composition information for patientspecific dose estimates.38

Dose for Other Breast Screening Modalities Because sensitivity for cancer detection is reduced in women with mammographically dense breasts (high glandular composition), several other breast imaging modalities are now under consideration for use as supplemental screening methods, including ultrasound, digital breast tomosynthesis (DBT), molecular breast imaging (MBI), and magnetic resonance imaging. Of these, only DBT and MBI use ionizing radiation. DBT is an add-on to conventional mammography and produces a limited angle 3D reconstruction of the breast. Keeping in mind the MQSA limit of 3 mGy on AGD to the 50-50 composition phantom, manufacturers have strived to design systems such that AGD from the combination of conventional FFDM combined with DBT would remain under this limit. AGD from DBT is estimated to be about equivalent to that of FFDM, such that if FFDM and DBT are used in combination, the AGD and effective dose are approximately doubled relative to FFDM alone.39 MBI is a nuclear medicine test that uses a dedicated gamma camera to image functional uptake of a radiopharmaceutical, typically Tc-99m sestamibi. Recent work has demonstrated feasibility of performing MBI using reduced administered activity of less than 300 MBq (8 mCi) with a dual-head cadmium zinc telluride detector.40,41 Intravenous administration of sestamibi delivers a systemic radiation dose to the body with varying absorbed dose among each organ.42 The dose to the breast from 300-MBq sestamibi is approximately 1.1 mGy, less than that reported for bilateral 2-view FFDM (4.7 mGy).37 Taking into account doses to all organs and their relative radiosensitivites, the corresponding effective dose from 300 MBq sestamibi is approximately 2.6 mSv,42 which is below the annual radiation dose from natural background and thus may be considered acceptable to use in a screening setting.

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Dose-Reduction Techniques Emerging Dose-Reduction Methods in CT The first step for radiation dose reduction in screening examinations is justification. Guidelines regarding the cohort that is appropriate for screening examinations (eg, age and other risk factors) can be found in societies’ recommendations and reports. Performing screening examinations only on these patients ensures that the benefit of these examinations outweighs potential risks. Having determined then that an examination is justified, the next step is to optimize scanning protocols using all available dose-reduction techniques to achieve sufficient image quality at low doses. During the past years, technology has been improved to continuously lower radiation dose in medical examinations.16,43,44 Scanners are getting more dose efficient. For example, detectors with integrated electronics were introduced to reduce electronic noise and consequently reduce radiation dose,45 and dynamic collimators that open and close at the 2 ends of the scan range reduce unnecessary dose at the periphery of a helical scan range.46 In this section, we review a few key techniques that have substantial effect on dose reduction in screening examinations. Automatic exposure control (AEC), or tube current modulation, is a technique that modulates tube current during the scan according to patient attenuation, which can vary dramatically over the scan range.47,48 The tube current can be modulated with the projection angle to account for attenuation differences in a scan plane (ie, lateral vs anterior-posterior direction), along the longitudinal direction to account for attenuation differences in different body regions (ie, chest vs abdomen), or with both angle and z-axis position (Figure 5). In general, tube current increases in regions with higher attenuation and decreases in regions with lower attenuation, although specific implementations vary among vendors. The aim was to redistribute the delivered photons to achieve the lowest dose for a specific image quality. AEC was not widely available at the

Figure 5 Diagram shows that CT scanners with automatic exposure control adjust the tube current (mA) according to patient attenuation. The high-frequency oscillations represent modulation of the tube current as the tube movies around the patient. The midpoint of the oscillations changes according to overall patient attenuation along the z-axis. (Color version of figure is available online.)

155 beginning of the NLST trial; therefore, its effect on dose reduction was not reflected in the reported dose estimates.20 Now, AEC is widely used in almost all diagnostic and screening examinations. Recently, there have been studies exploring the effect of tube potential (kV) on image quality and radiation dose. Images acquired at lower kV have higher signal for high z materials, such as iodine and calcium. This effect has the potential to reduce radiation dose if the same iodine or calcium signal is maintained. However, the lower kV images are also usually associated with higher image noise due to lower penetration capability. Therefore, dose-reduction potential depends on both imaging task and patient size. A method of automatically selecting optimal kV based on the imaging task and patient size has been developed, and studies have demonstrated 25%-36% dose reduction in iodine contrast-enhanced abdominopelvic and CT angiography examinations.49,50 For CAC screening, calcium signal changes with respect to kV. Therefore, potential dose reduction using lower kV is possible, even though the examination is a noncontrast scan. A study by Marwan et al51 demonstrated that CAC with 100 kV reduced dose in a highpitch scan mode compared with 120 kV. Lung cancer and CTC screening are noncontrast scans, and nodules and polyps are tissuelike materials that do not gain substantial contrast by changing kV. Dose reduction for these scans by changing kV is limited. Using chest phantoms, Yu et al demonstrated some dose reduction in noncontrast chest CT scans for small pediatric patients, but no dose reduction for regular and large size adults. Importantly, lower kV might not always be the optimal kV; 140 kV can be the optimal kV for large size patients. Beam filtration plays a key role in CT imaging. The bow-tie or beam-shaping filter is used to reduce the dose to peripheral body regions. A dedicated cardiac bow-tie filter has been shown to reduce radiation dose in cardiac scans.52 Besides bow-tie filters, added flat filters have been recently introduced to change the beam spectrum to achieve lower radiation dose in certain examinations, such as lung cancer screening and colon cancer screening.53 The added filter removes most of the lower energy x-ray photons and shifts the beam spectrum to higher energies. As most of the lower energy photons are absorbed by the patients, without contributing to image formation, they contribute to unnecessary patient dose. Ultra–low-dose chest CT with 0.06-mSv effective dose has been reported using an added flat tin filter.53 Recently, iterative reconstruction (IR) algorithms have been made commercially available to reduce the radiation doses from CT.54-58 Compared with conventional filtered backprojection (FBP) algorithms, IR has the advantage of accurately modeling photon statistics and system geometry, which can produce lower image noise than conventional FBP algorithms, given the same radiation dose. Dose reduction is therefore possible while maintaining the same image noise. Figure 6 illustrates the IR process. From the measured projection data, an initial guess (ie, image reconstructed with FBP algorithm) is made and then simulated projections are calculated from the initial guess image. Simulated projection data are then compared with measured projection data, and

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156 the differences are used to correct the initial guess images. This procedure is repeated until conversion criteria are met. All major CT vendors now have IR solutions available on commercial CT scanners. These IR algorithms can be image based, projector data based, or hybrid based. Image-based IR iterate only in the image space to reduce image noise, whereas projection data–based IR goes back and forth between image and projection space. Image-based methods are much faster than projection data–based ones, whereas projection data– based methods can potentially produce better image quality with fewer image artifacts. Hybrid approaches try to maintain the advantages of both approaches. Substantial dose reduction using IR has been demonstrated in the screening examinations of CAC,59,60 lung cancer,61-64 and CTC.65-67 Noise-reduction techniques other than IR have also been investigated. These algorithms reduce image noise by filtering either in the projection space or in the image space. A key feature of the filter design is that it maintains edge details while reducing image noise. Multiple algorithms have been developed, such as bilateral filters, nonlocal means filters, and 3D diffusion filters.68-70 For older CT scanners without IR, these noise reduction techniques can be an economical solution for dose reduction. Given the ECG-gated acquisition of cardiac CT, unique dose-reduction techniques have been developed. A major dose-reduction technique is ECG-based tube current modulation, in which the tube current is lowered to 4%-20% of the regular values during portions of the cardiac phase where coronary artery visualization is not required. This results in a substantial dose reduction for retrospectively gated cardiac scans. Up to 64% dose reduction was found in a coronary CT angiography study.71 Generally speaking, prospectively gated scans are associated with lower radiation dose compared with retrospectively gated scans, whereas retrospectively gated scans are more capable of dealing with heart rate variation and irregular heartbeats and allow functional analysis. Several newly introduced techniques have enabled further dose reduction in CAC and cardiac CT, including use of a large area detector (up to 16 cm along the z-axis) to cover the whole heart in 1 gantry rotation, and a high-pitch mode that enables helical scans to be acquired in 1 heart beat.51,72,73 Besides the advanced technologies mentioned earlier, there are also practical procedures and steps that should be taken to

Figure 6 The process of iterative reconstruction algorithms in CT. (Color version of figure is available online.)

minimize radiation dose. For example, the scanning range should cover only the anatomy of interest, as scan range is directly related to the integrated dose the patient receives. This simple yet effective method can reduce unnecessary radiation to the patient without affecting diagnostic performance. Scanning protocols should be regularly reviewed and optimized to incorporate newly developed dose-reduction techniques.

Recent Dose-Reduction Methods in Mammography Traditionally, mammographic x-ray tubes have employed molybdenum (Mo) targets with Mo filters, or for thicker breasts, rhodium (Rh) filters, to produce x-rays with an average energy of about 20 keV. This low energy is necessary for providing optimal contrast of soft tissues in the breast. Some digital mammographic systems now use a tungsten (W) anode with Rh filters and, for thicker breasts, silver (Ag) filters, yielding an x-ray spectrum with higher energies that results in reduced AGD.74 One study reported AGD was reduced by approximately 50% with the W/Rh system: average AGD in patient examinations was 2.3 mGy with Mo-Mo, 2.8 mGy with Mo-Rh, and 1.3 mGy with W-Rh.75 Another study reported average AGD of approximately 1.5 mGy for Mo-Rh vs 1 mGy with W-Rh.76 The recent introduction of mammographic systems comprising a slot-scanning collimator coupled with a photoncounting detector has resulted in the lowest AGDs reported thus far, while maintaining excellent diagnostic performance. In a retrospective analysis of women undergoing screening in Germany with either conventional FFDM or a photon-counting system, AGD was approximately 0.6 mGy for photon-counting, which was significantly lower than 1.7 mGy for conventional FFDM, and a higher cancer detection rate was obtained with the photon-counting systems.77

CT Dose Check and Dose Registries With the increased attention to the radiation doses used in medical imaging, efforts have been invested in monitoring and controlling radiation dose. A National Electrical Manufacturers Association standard was recently published, which is referred to as “Dose Check.”78 Dose Check is a tool to notify users before CT scanning when the dose might be too high. This allows the user to confirm or correct a scanning technique that might lead to an inadvertent high dose. The Dose Check tool uses 2 sets of values: dose notification and dose alert values. Dose notification value is the dose that triggers a notification because the dose value would likely be exceeded by the prescribed scan. This can be programmed for individual protocols and individual scans. The dose alert value is a global setting for all protocols. It is the dose that triggers an alert because the total dose in an ongoing examination results in a cumulative dose that exceeded the user-configured alert value.

Ionizing radiation in screening examinations The American Association of Physicists in Medicine (AAPM) has provided recommendations for dose notification and alert values.79 The effect of this standard on clinical workflow has been investigated.80 A correctly performed screening examination should never trigger a dose alert event. Because of the low-dose nature of screening examinations, notification values for screening examinations should be lowered compared with those of diagnostic examinations. There is also a strong interest in comparing doses between practices. The ACR Dose Index Registry collects CTDIvol and other critical data elements from anonymized patient data submitted by more than 800 facilities. This allows facilities to compare their CT dose indices to regional and national values. Institutions are provided with periodic feedback reports comparing their results, by body part and examination type, to aggregate results.81 Commercial and home-grown computer software has also been developed to track patient dose information. This dose information can be obtained from the Digital Imaging and Communications in Medicine header of the patient images, the patient protocol page, or the dosestructured report. This information can be used for quality control and practice improvement.

Radiation Risks From Screening Tests No studies have demonstrated direct carcinogenic effects from low-dose radiation at the levels of medical imaging examinations, with effective doses of approximately less than 10 mSv. Carcinogenic effects have been observed from radiation at much higher doses (effective dose exceeding 100-150 mSv), as has been shown in follow-up studies of atomic bomb survivors and patients undergoing high-dose radiation treatments.82,83 Less than 100 mSv, carcinogenic effects are extremely difficult to prove or disprove. Hence, there is an ongoing debate on how to assign risk for effective doses o100 mSv. A number of studies have dealt with this problem by assuming a linear-no-threshold (LNT) model for the relationship between radiation and risk, that is, the association of cancer risk and radiation found at high doses can be linearly extrapolated down to very low levels, and there is no threshold below which no risk exists. Using this LNT model, studies have made hypothetical calculations of cancer incidence and mortality for radiation doses in the diagnostic imaging range.14,84–86 These studies estimated risks using data published in the BEIR report, accounting for age and sex of the patient. However, a number of important uncertainties are present in such calculations that cast doubt on the association of cancer risk and low-dose radiation.87 Most relevant, the LNT model is likely inappropriate for use at low doses, as suggested by a large collection of radiobiological and human exposure data. Additionally, the Japanese atomic bomb survivors, from which much of the risk estimates are based, were exposed to a substantially higher dose rate (instantaneous blast) and substantially different type of radiation (high-energy gamma-rays, neutrons, and charged particles) than patients undergoing

157 Table 2 Typical Doses for Common Screening Examinations as Well as Other Sources of Radiation Examination

Reported Effective Dose (mSv)

CT coronary artery calcium screening

0.6-1.7 For prospective cardiac gating91 2.5-5.3 For retrospective cardiac gating91 CT lung cancer screening Mean ¼ 1.4 (range: 0.5-4)20 CT colonography screening Median ¼ 5.7 (range: 2.612.2)25 Film screen mammography Mean ¼ 0.60 (range: 0.510.64)37 Full-field digital Mean ¼ 0.50 (range: 0.45mammography 0.60)37 Digital breast tomosynthesis 0.539 Molecular breast imaging 2.6 (for 8 mCi Tc-99m sestamibi)42 Annual background radiation, Mean ¼ 3 (range: 1-10) worldwide Chest x-ray Mean ¼ 0.05 (range: 0.0160.3)

medical imaging. Lastly, a large number of uncertainties exist within the BEIR report, as described within its contents, and many of the risk coefficients and weighting factors were determined by committee opinion, not absolute evidence.88 Considering all available evidence regarding radiation doses and risk, scientific and professional societies, such as the AAPM, Health Physics Society, and United Nations Scientific Council on the Effects of Atomic Radiation, composed of radiation experts have released position statements regarding the practice of estimating risk from medical imaging tests that use ionizing radiation. Both the AAPM and the Health Physics Society recommend against the estimation of health risks for single procedures less than effective doses of 50 mSv, or for multiple procedures that accumulate to less than 100 mSv. Both societies state that health effects less than 50 mSv, which is also the annual limit for radiation workers in the United States, are too low to be observed and may not exist. United Nations Scientific Council on the Effects of Atomic Radiation has specifically addressed the large number of uncertainties in extrapolation of high-dose risk data to low doses, in other populations and for different types of radiation, and ultimately has recommended against the practice of “multiplying [risks from] low doses by large numbers of individuals to estimate numbers of radiation-induced health effects within a population exposed to incremental doses at levels equivalent to or below natural [annual] background levels.”89,90 The mean annual background effective dose in the United States is 3 mSv, with a range of 1-10 mSv. The radiation doses from the screening examinations discussed here (CT screening for coronary artery calcium levels, lung and colon cancers, mammography, and tomosynthesis or MBI for breast cancer) are all less than or within the lower half of this range (Table 2). Thus, when screening examinations have been demonstrated to provide health benefits in target populations, they should be

158 recommended and performed; the risk of negative, long-term health effects from radiation is either too small to be observed or does not exist.

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