Mammography Acquisition: Screen-Film, Digital Mammography and Tomosynthesis, the Mammography Quality Standards Act, and Computer-Aided Detection

Screen-Film Mammography Image Acquisition

In SFM, the image receptor assembly holds a screen-film cassette in a carbon-fiber support with a moving antiscatter grid in front of the cassette and an AEC detector behind it (see Fig. 1.3A ). Screen-film image receptors are required to be 18 × 24 cm and 24 × 30 cm in size to accommodate both smaller and larger breasts ( Box 1.5 ). Each size image receptor must have a moving antiscatter grid composed of lead strips with a grid ratio (defined as the ratio of the lead strip height to the distance between strips) between 3.5:1 and 5:1. The reciprocating grid moves back and forth in the direction perpendicular to the grid lines during the radiographic exposure to eliminate grid lines in the image by blurring them out. One manufacturer uses a hexagonal-shaped grid pattern to improve scatter rejection; this grid is also blurred by reciprocation during exposure. Use of a grid improves image contrast by decreasing the fraction of scattered radiation reaching the image receptor. Grids increase the required exposure to the breast by approximately a factor of 2 (the Bucky factor), because of the attenuation of primary, as well as scattered, radiation. Grids are not used with magnification mammography. Instead, in magnification mammography, scatter is reduced by collimation and by rejection of scattered x-rays due to a significant air gap between the breast and image receptor.

The AEC system, also known as the phototimer , is calibrated to produce a consistent film optical density (OD) by sampling the x-ray beam after it has passed through the breast support, grid, and cassette. The AEC detector is usually a D-shaped sensor that lies along the midline of the breast support and can be positioned by the technologist closer to or farther from the chest wall. If the breast is extremely thick or inappropriate technique factors are selected, the AEC will terminate exposure at a specific backup time (usually 4–6 seconds or 300–750 mAs) to prevent tube overload or melting of the x-ray track on the anode.

Screen-film cassettes used in mammography have an inherent spatial resolution of 18 to 21 lp/mm. Such resolution is achieved typically by using a single-emulsion film placed emulsion side down against a single intensifying screen that faces upward toward the breast in the film cassette. The single-emulsion film with a single intensifying screen is used to prevent the parallax unsharpness and crossover exposure that occur with double-emulsion films and double-screen systems. One manufacturer has introduced a double-emulsion film with double-sided screens (EV System, Carestream Health, formerly Eastman Kodak Health Group) with a thinner film emulsion and screen on top to minimize parallax unsharpness. Most screen-film processing combinations have relative speeds of 150 to 200, with speed defined as the reciprocal of the x-ray exposure (in units of Roentgen) required to produce an OD of 1.0 above base plus fog (1.15–1.2, because base plus fog OD is 0.15–0.2).

Film processing involves development of the latent image on the exposed film emulsion. The film is placed in an automatic processor that takes the exposed film and rolls it through liquid developer to amplify the latent image on the film, reducing the silver ions in the x-ray film emulsion to metallic silver, resulting in film darkening in exposed areas. The developer temperature ranges from 92°F to 96°F. The film is then run through a fixer solution containing thiosulfate (or hypo ) to remove any unused silver and preserve the film. The film is then washed with water to remove residual fixer, which if not removed can cause the film to turn brown over time. The film is then dried with heated air.

Film processing is affected by many variables, and the most important is developer chemistry (weak or oxidized chemistry makes films lighter and lower contrast), developer temperature (too hot may make films darker, and too cool may make films lighter), developer replenishment (too little results in lighter, lower contrast films), inadequate agitation of developer, and uneven application of developer to films (causing film mottling; Table 1.2 ).

Film viewing conditions must be appropriate ( Fig. 1.4 ). Because mammography viewboxes have high luminance levels (>3000 cd/m ² [3000 nit]), mammograms should be masked so that no light strikes the radiologist’s eye without passing through the exposed film. Because of high luminance levels film collimation of x-ray exposure should be rectangular and extend slightly beyond the edge of the image receptor so that film is darkened to its edges. Viewbox luminance should be reasonably uniform across all viewbox panels. In addition, the ambient room illumination should be low (<50 lux, and preferably less) to minimize “dazzle glare” from film surfaces. Both viewbox luminance and room illumination should be checked annually by the medical physicist as part of the site quality control program, as specified in the ACR Mammography Quality Control Manual .

Digital Mammography Image Acquisition

In digital mammography, the image is obtained in the same manner as in screen-film mammography, using a compression plate and an x-ray tube, with the screen-film cassette replaced by a digital detector (see Fig. 1.3B ). Digital image acquisition has several potential advantages in terms of image availability, image processing, making annotations ( Fig. 1.5 ), and CAD. One advantage is elimination of the film processor, which eliminates artifacts and image noise added during film processing. The image contrast of digital mammography is different among vendors depending on the digital look-up curve, which governs how digital signals are translated into pixel gray scale values. Figure 1.6 shows digital mammograms that were obtained with two machines from different vendors demonstrating how the image contrast varies.

Digital mammography uses indirect or direct digital detectors. Indirect digital detectors use a fluorescent screen made of materials such as cesium iodide (CsI) to convert each absorbed x-ray to hundreds of visible light photons. Behind the fluorescent material, light-sensitive detector arrays made of materials such as amorphous silicon diodes or charge-coupled devices measure the produced light pixel by pixel. The weak electronic signal measured in each pixel is amplified and sent through an analog-to-digital converter, enabling computer storage of each pixel’s measured detector signal.

Direct digital detectors use detector elements that capture and count x-rays directly, although amplification and analog-to-digital conversion are still applied. Another method to produce digital mammograms involves amorphous selenium. An amorphous selenium plate is an excellent absorber of x-rays and an excellent capacitor, storing the charge created by ionization when x-rays are absorbed. After exposure, an electronic device is used to read out the charge distribution on the selenium plate, which is in proportion to local exposure. This can be done by scanning the selenium plate with a laser beam or by placing a silicon diode array in contact with one side of the plate, with bias voltage applied, to read out the stored charge. Each of these methods allows production of high-resolution digital images.

Another approach to full-field digital mammography (FFDM) is computed radiography (CR), which uses a photostimulable phosphor composed of barium fluorobromide doped with europium (BaFBr:Eu). Computed radiography uses the same dedicated mammography units as SFM, replacing the screen-film cassettes and film processor with CR cassettes (in sizes of 18 × 24 cm and 24 × 30 cm) and a CR processor. The phosphor plate within the CR cassette is used to absorb x-rays just as the screen in a screen-film cassette. Rather than emitting light immediately after exposure (through fluorescence), x-ray absorption in the phosphor causes electrons within the phosphor crystals to be promoted to higher energy levels (through photostimulation). The plate is removed from the cassette in the CR processor and a red laser light scans the phosphor plate point by point, releasing electrons and stimulating emission of a higher energy (blue) light in proportion to x-ray exposure. In conventional x-ray systems, CR phosphor plates have an opaque backing and are read from only one side. In at least one FDA-approved CR system for mammography (Fuji 5000D CR, Fujifilm Medical Systems), the CR cassette base is transparent and light emitted from the plate during laser scanning is read from both sides to increase reading efficiency.

No matter which digital detector is used, its job is to measure the quantity of x-rays passing through the breast, compression plate, grid (in contact mammography), and breast holder. The signal measured in each pixel is determined by the total attenuation in the breast along a given ray.

The choice of an analog-to-digital converter determines how many bits of memory will be used to store the signal for each pixel; the more bits per pixel, the more dynamic range there is for the image, but at a higher digital data storage cost. Specifically, if 12 bits per pixel are used, 2 ¹² or 4096 signal values can be stored. If 14 bits per pixel are used, 2 ¹⁴ or 16,384 signal values can be stored. Usually 12- to 14-bit storage per pixel is used. In either case, 2 bytes per pixel are required (8 bits = 1 byte) to store the image. For example, the GE Senographe 2000D and DS digital detectors have 1920 × 2304 pixel arrays, or 4.4 million pixels, requiring 8.8 million bytes (8.8 megabytes; MB) of storage per image. Other FFDM systems require up to 52 MB of storage per image.

Screen-film image receptors used for mammography have a line pair resolution of 18 to 21 lp/mm. To equal this spatial resolution, a digital detector would require 25-micron pixels, which would yield noisier images and pose a storage issue caused by the large data sets required to store those images. FFDM systems have spatial resolutions ranging from 5 lp/mm (for 100-micron pixels) to 10 lp/mm (for 50-micron pixels). In digital mammography systems, it is the size of the pixels, or more correctly their center-to-center distance (pitch), that determines (and limits) the spatial resolution of the imaging chain.

The lower limiting spatial resolution of FFDM systems compared with film is offset by the increased contrast resolution of FFDM systems. Unlike SFM, in which the image cannot be manipulated after exposure and processing, FFDM images can be optimized after image capture by image postprocessing and adjustment of image display. For fixed digital detectors, such as CsI and silicon diode arrays (used by GE) and selenium and amorphous silicon diode arrays (used by Hologic and Siemens), one image-processing step that can minimize image noise and structured artifacts is flat-field correction, or gain correction of each acquired digital image. This is done by making and storing a sensitivity map of the digital detector and using that map to correct all exposures. Typically, slot-scanning devices (such as the older SenoScan digital system, Fischer Medical Systems) and CR systems do not perform flat-field correction of digital images. Beyond this, all digital systems have the ability to process the acquired digital image to minimize or eliminate the signal difference that results from the roll-off in thickness of the breast toward the skin line (thickness equalization); some devices add processing to help enhance the appearance of microcalcifications (eg, GE Premium View and FineView). The window width and window level for all digital images viewed with soft copy display on review workstations can be adjusted, changing the contrast and brightness of the images, respectively, as well as digitally magnifying images.

Another important difference between SFM and FFDM is that screen-film images have a linear relationship between the logarithm of x-ray exposure and film OD only in the central portion of the characteristic curve. In FFDM, there is a linear relationship between x-ray exposure and signal over the entire dynamic range of the detector. Thus digital images (at least their “raw” or “for processing” presentation) do not suffer contrast loss in underexposed or overexposed areas of the mammogram (as long as detector saturation does not occur); instead, they show similar contrast over the full dynamic range of signals. Different manufacturers apply different look-up tables to digital images in transforming them from initially acquired raw or for processing images to processed or for presentation images. These different look-up tables affect the contrast of final presentation digital images. Some, such as Hologic’s linear look-up table, yield higher contrast images, whereas others, such as GE’s sigmoidal look-up table, yield images presented with less contrast and more like screen-film images. In either case, thickness equalization is used to equalize signal differences from the center of the breast to the skin line. FFDM also has the advantage of eliminating the variability and noise added by film processing that is inherent to SFM.

In terms of breast dose, FFDM has a mean glandular dose lower than, or comparable with, the radiation dose of SFM. Results from the American College of Radiology Imaging Network (ACRIN) DMIST found the average single-view mean glandular dose for FFDM to be 1.86 mGy, 22% lower than the average SFM mean glandular dose of 2.37 mGy ( ). Specific manufacturers, especially those using slot-scanning techniques, produce lower doses than SFM. Slot-scanning systems have a narrow slot of detector elements that are scanned under the breast in synchronization with a narrow fan beam of x-rays swept across the breast. This design, although more technically difficult to implement, has the advantage of eliminating the need for a grid to reduce scattered radiation. Scatter is partially eliminated by the narrow slot itself. The absence of a grid reduces the amount of radiation to the breast needed to get the same SNR in the detector. Most full area digital detectors also have demonstrated lower breast doses compared with SFM, especially for thicker breasts.

Once captured and processed, the image data are transferred to a reading station for interpretation on high-resolution (2048 × 2560 or 5-Mpixel) monitors or printed on film by laser imagers (with approximately 40-micron spot sizes, so that film printing does not reduce the inherent spatial resolution of digital mammograms) for interpretation of hardcopy images on film viewboxes or alternators ( Fig. 1.7 ). Digital data can be stored on optical disks, magnetic tapes, picture archiving and communication systems (PACS), or CDs for later retrieval.

The MQSA states that FFDM images must be made available to patients as hardcopy films, as needed, which means the facility must have access to an FDA-approved laser printer for mammography that can reproduce the gray scale and spatial resolution of FFDM images. The images may also be given to the patient on a CD with an image viewer, if this is acceptable to the patient.

A number of studies have evaluated the performance of FFDM compared with SFM for screening asymptomatic women for breast cancer. Early studies showed comparable or slightly worse results (but not statistically significant differences) for receiver operating characteristic (ROC) curve area and sensitivity ( ) or cancer detection rate ( ) of FFDM compared with SFM. Larger studies, however, showed some benefits of FFDM compared with SFM. The ACRIN DMIST paired study ( ) showed no difference overall, but found that FFDM had statistically significantly higher ROC curve areas than SFM for women under age 50, for premenopausal and perimenopausal women, and for women with denser breasts (Breast Imaging Reporting and Data System [BI-RADS] density categories C and D). These findings are supported by who showed in clinical practice in the United States that FFDM had higher, but not necessarily significantly higher, sensitivity than SFM in most age groups, including women 40 to 49, premenopausal and perimenopausal women, and women with extremely dense breasts. The sensitivity of FFDM was significantly higher than for SFM among women aged 40 to 79 who had estrogen receptor–negative cancers, and especially so among women aged 40 to 49 (95% versus 55%; p = 0.007) The Oslo II trial showed that digital mammography had a significantly higher cancer detection rate (5.9 cancers per 1000 women screened) than SFM (3.8 cancers per 1000 women screened) ( ).

Interpretation times for screening exams using FFDM tend to take 1.5 to 2 times longer than screening exams on SFM ( ; ). As of May 2016, 98% of the mammography units in the United States are digital mammography systems. The FDA-approved manufacturers for digital mammography units and their approval dates are listed in Box 1.6 .

Tomosynthesis Acquisition

Digital breast tomosynthesis obtains a set of rapidly acquired low-dose digital projections taken at multiple angles through the compressed breast to reconstruct a stack of high-resolution, mammographic-quality planar images through the entire breast ( Fig. 1.8 ). The reconstructed planes are parallel to the plane of the breast support (perpendicular to the central ray of the x-ray unit) and are spaced every 0.5 mm or 1 mm apart. The technique is similar to the previous technique of linear film tomography, but with one major difference. In film tomography, a full sweep of the x-ray tube resulted in a single planar image through the patient, with tissues in the selected plane in focus while tissues outside that single plane were blurred. In digital tomosynthesis, a single sweep of the x-ray tube and reconstruction of the stored digital data results in a stack of dozens of parallel images through the breast, each image with a single in-focus plane, with blurring of the tissues above and below each focal plane.

The stack of reconstructed in-focus planar images from DBT permits improved visualization of lesion margins and can reveal suspicious lesions with greater clarity than conventional mammography ( Fig. 1.9 ) ( ). This is possible because DBT minimizes structured noise caused by overlapping tissues, which is a significant limitation of conventional two-dimensional (2D) screen-film or digital mammography. DBT images also reduce callbacks for additional imaging by eliminating or reducing the complication of superimposed fibroglandular tissues that can appear suspicious in conventional 2D projection mammography.

Different manufacturers have taken different approaches to DBT in terms of the number of acquired low-dose projections, angular range of the projections, detector types, and scan time. Table 1.3 presents DBT design parameters for four different DBT manufacturers. As of May 2016, three manufacturers (Hologic, GE Healthcare, and Siemens Healthcare) have received FDA approval for clinical use of DBT for screening and diagnostic breast imaging in the United States.

A multireader FDA approval study of Hologic DBT by found significantly higher ROC curve areas for two-view DBT added to two-view FFDM compared with FFDM alone. Most clinical studies of DBT have reported similar results, with higher breast cancer detection rates and/or higher sensitivity with two-view DBT added to two-view FFDM, than with FFDM alone. They have also shown significantly lower recall rates with DBT plus FFDM compared with DBT alone. For example, compared the performance of 13,856 screening FFDM studies finding 56 cancers in 2010 to 9499 studies with FFDM plus DBT finding 51 cancers from May 2011 to early 2012. They found that cancer detection rates increased from 4.0 to 5.4 per 1000 screenings ( p = 0.18, not significant), whereas recall rates dropped from 8.7% to 5.5% ( p < 0.001, highly significant), and positive predictive value for recalls increased from 4.7% to 10.1% ( p < 0.001). Similar results were found by in Oslo and by in Italy: significantly higher cancer detection rates and significantly lower false-positive rates with DBT plus FFDM compared with FFDM alone.

Breast radiation doses with DBT added to FFDM are approximately double that of FFDM alone ( ). Recently, however, Hologic has received FDA approval to replace DBT plus FFDM acquisitions with DBT acquisitions in which 2D FFDM (C view) images are synthesized from DBT data, reducing breast radiation doses from two-view DBT with C view images to approximately the same as two-view FFDM. GE’s approach to DBT approved by the FDA was to acquire a DBT mediolateral oblique (MLO) view and 2D craniocaudal (CC) FFDM view of each breast at the same dose as two-view FFDM. GE systems also have the capability to acquire DBT CC views at approximately the same dose as 2D FFDM CC views.

Interpretation times for screening exams using DBT plus FFDM tend to be longer than for FFDM alone by 47% to 102% ( ; ).

Views and Positioning

The two views obtained for screening mammography are the CC and MLO projections. The names for the mammographic views and abbreviations are based on the ACR BI-RADS, a lexicon system developed by experts for standard mammographic terminology. The first word in the mammographic view indicates the location of the x-ray tube, and the second word indicates the location of the image receptor. Thus a CC view would be taken with the x-ray tube pointing at the breast from the head (cranial) down through the breast to the image receptor in a more caudal (toward the feet) position.

For positioning, the technologist tailors the mammogram to the individual woman’s body habitus to get the best image. The breast is relatively fixed in its medial borders near the sternum and the upper breast, whereas the lower and outer portions of the breast are more mobile. The technologist takes advantage of the mobile lower outer breast to obtain as much breast tissue on the mammogram as possible. One component of ACR Mammography Accreditation is clinical image review in which clinical images are submitted for peer review of one patient with dense breasts and one patient with fatty breasts acquired on each mammography unit every 3 years. To pass ACR accreditation clinical image review, the MLO mammogram must show most of the breast tissue in one projection, with portions of the upper inner and lower inner quadrants partially excluded ( Figs. 1.10 and 1.11 ). Clinical evaluation of the MLO view should show fat posterior to the fibroglandular tissue and a large portion of the pectoralis muscle, which should be convex and extend inferior to the posterior nipple line (PNL; Figs. 1.12 and 1.13 ). The nipple must be in profile on at least one of the two images. The PNL describes an imaginary line drawn from the nipple to the pectoralis muscle or chest wall image edge and perpendicular to the pectoralis muscle. The PNL should intersect the pectoralis muscle in the MLO view in more than 80% of women. The MLO view should show adequate compression, exposure, contrast, and an open inframammary fold ( Fig. 1.14 ), in which both the lower portion of the breast and a portion of the upper abdominal wall should be seen.

To pass ACR clinical image review, the CC view should include the medial posterior portions of the breast without sacrificing the outer portions ( Fig. 1.15 ; see Fig. 1.10 ). With proper positioning technique, the technologist should be able to include the medial portion of the breast without rotating the patient medially by lifting the lower medial breast tissue onto the image receptor. The pectoralis muscle should be seen when possible on the CC view. On the CC view, the PNL extends from the nipple to the pectoralis muscle or the chest wall edge of the image and perpendicular to the pectoralis muscle or image edge. For a given breast, the length of the PNL on the CC view should be within 1 cm of its length on the MLO view.

Although the technologist tries to avoid producing skin folds on the image when possible, they are seen occasionally and sometimes the image needs to be repeated, and other times they may not cause problems for the radiologist reading the image ( Fig. 1.16 ).

described reasons why 1034 mammographic clinical images failed ACR accreditation, which included positioning in 20%, contrast in 13%, labeling in 8%, and noise in 5% at a time when only film-screen mammography was available. They included the deficiencies mentioned previously as reasons for failure and also included these reasons: sagging breast tissue, portions of the breast not visualized, other body parts included on the mammogram, breast positioned too high on the image receptor, motion artifact, poor compression resulting in poor separation of breast tissues, poor contrast or exposure, and unsharpness.