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Fig. 1-8. An early Lunar DP3 dual-photon absorptiometer. This device utilized 153Gd to generate photon energy. Photo courtesy of GE Medical Systems, Madison, WI.

DPA studies of the spine required approximately 30 minutes to complete. Studies of the proximal femur took 30 to 45 minutes to perform. Total body bone density studies with DPA required 1 hour. Skin radiation dose was low during spine or proximal femur studies at 15 mrem. Accuracy of DPA measurements of the spine ranged from 3 to 6% and for the proximal femur, 3 to 4% (48). Precision for measurements of spine bone density was 2-4% and around 4% for the femoral neck.

DPA was considered a major advance from SPA because it allowed the quantification of bone density in the spine and proximal femur. DPA did have several limitations, however. Machine maintenance was expensive. The 153Gd source had to be replaced yearly at a cost of $5000 or more. It had also been noted that as the radioactive source decayed, values obtained with DPA increased by as much as 0.6% per month (49). With replacement of the source, values could fall by as much as 6.2%. Although mathematical formulas were developed to compensate for the effect of source decay, it remained a cause for concern, potentially affecting both accuracy and precision. The precision of 24% for DPA measurements of the spine and proximal femur limited its application in detecting changes in bone density. With a precision of 2%, a change of at least 5.5% from the baseline value had to be seen before one could be certain at the 95% confidence level that any change had occurred at all (50). With a precision of 4%, this figure increased to 11.1%. At a lower 80% confidence level, the required changed for precision values of 2 and 4% were 3.6 and 7.2%, respectively. As a practical matter, this meant that DPA bone density studies would not show significant changes for up to 5 years. This was too long a period to wait to be clinically useful.

In DPA spine bone density studies in which the photon beam passed in a PA direction, the highly trabecular vertebral body could not be separated from its more cortical posterior elements. In addition, the cortical shell of the vertebral body could not be separated from its trabecular interior. Calcifications in the overlying soft tissue or abdominal aorta will attenuate such a beam, falsely elevating the bone density values. Arthritic changes in the posterior elements of the spine also affect the measurement (51). These effects are discussed in greater detail in Chapters 2 and 9. PA DXA studies of the spine are not immune to these effects either but lateral DXA spine studies can be performed to overcome these limitations. Studies of the spine in the lateral projection were never available with DPA.

The ability to make site-specific predictions of fracture risk of the spine and proximal femur or global fracture risk predictions with DPA was established in prospective trials (19,39). Like SPA, DPA is rarely performed in the United States now, because of the availability of DXA with its technological improvements.

Dual-Energy X-Ray Absorptiometry

The underlying principles of DXA are the same as those of DPA. With DXA, however, the radioactive isotope source of photon energy has been replaced by an X-ray tube. There are several advantages of X-ray sources over radioactive isotopes. There is no source decay that would otherwise require costly replacement of the radioactive source. Similarly, there is no concern of a drift in patient values due to source decay. The greater source intensity or "photon flux" produced by the X-ray tube and the smaller focal spot allows for better beam collimation resulting in less dose overlap between scan lines and greater image resolution. Scan times are faster and precision is improved.

Because X-ray tubes produce a beam that spans a wide range of photon energies, the beam must be narrowed in some fashion in order to produce the two distinct photoelectric peaks necessary to separate bone from soft tissue. The major manufacturers of central DXA systems in the United States have chosen to do this in one of two ways. GE Medical Systems of Madison, WI and Norland, a CooperSurgical Company of Fort Atkinson, WI, use rare earth K-edge filters to produce two distinct photoelectric peaks. Hologic Inc. of Bedford, MA uses a pulsed power source to the X-ray tube to create the same effect.

K-edge filters produce an X-ray beam with a high number of photons in a specific range. The energy range that is desired is the energy range that is just above the K-absorption edge of the tissue in question. The K-edge is the binding energy of the K-shell electron. This energy level varies from tissue to tissue. The importance of the K-edge is that at photon energies just above this level, the transmission of photons through the tissue in question drops dramatically. That is, the photons are maximally attenuated at this energy level (52). Therefore, to separate bone from soft tissue in a quantifiable fashion, the energy of the photon beam should be just above the K-edge of bone or soft tissue for maximum attenuation. GE Medical Systems uses a cerium filter in its central8 devices that has a K-shell absorption edge at 40 keV. A cerium-filtered X-ray spectrum at 80 kV will contain two photoelectric peaks at about 40 and 70 keV. The samarium K-edge filter employed by Norland in its central devices has a K-shell absorption edge of 46.8 keV. The samarium-filtered X-ray beam at 100 kV produces a low-energy peak at 46.8 keV. In the Norland system, the high-energy peak is variable because the system employs selectable levels of filtration but the photons are limited to less than 100 keV by the 100 kV employed. The K-edge of both cerium and samarium results in a low-energy peak that approximates the 44 keV low-energy peak of gadolinium-153 used in old dual-photon systems.

Hologic central DXA devices utilize a different system to produce the two photoelectric peaks necessary to separate bone from soft tissue. Instead of employing K-edge filtering of the X-ray beam, Hologic employs alternating pulses to the X-ray source at 70 kV and 140 kV.

Most regions of the skeleton are accessible with DXA. Studies can be made of the spine in both an posterior-anterior (PA)9 and lateral projection. Lumbar spine studies acquired in the lateral projection have the ability to eliminate the confounding effects of dystrophic calcification on densities measured in the PA direction (53). Lateral scans also eliminate the highly cortical posterior elements that contribute as much as 47% of the mineral content measured in the PA direction (54). The utility of lateral DXA lumbar spine studies can be limited by rib overlap of L1 and L2 and pelvic overlap of L4, more so when performed in the left lateral decubitus position than the supine position (53,55). Bone density in the proximal femur, forearm, calcaneus, and total body can also be measured with DXA.

Scan times are dramatically shorter with DXA compared to DPA. Early DXA units required approximately 4 minutes for studies of the PA lumbar spine or proximal femur. Total body studies required 20 minutes in the medium scan mode and only 10 minutes in the fast scan mode. Newer DXA units scan even faster, with studies of the PA spine or proximal femur requiring less than 1 minute to perform.

The values obtained with DXA studies of the skeleton are highly correlated with values from earlier studies performed with DPA. Consequently, the accuracy of DXA is consid

8 A central device is a bone densitometer that can be used to quantify bone density in the spine and proximal femur. The distinction between central and peripheral devices is discussed in Chapter 2.

9 Although spine bone density studies with DXA are often referred to as PA spine studies, the beam actually passes in a posterior to anterior direction. Such studies are correctly characterized as PA spine studies, but accepted convention is to refer to them as PA spine bone density studies. The Lunar Expert, a fan-array DXA scanner, does acquire spine bone density studies in the PA projection.

ered comparable to that of DPA (56-59). DXA spine values and Hologic and Norland DXA proximal femur values are consistently lower than values obtained previously with DPA. There are also differences in the values obtained with DXA equipment from the three major manufacturers.20 Values obtained with either a Hologic or Norland DXA unit are consistently lower than those obtained with a Lunar DXA unit, although all are highly correlated with each other (60-62). Comparison studies using all three manufacturers' central DXA devices have resulted in the development of formulas that make it possible to convert values for the lumbar spine and femoral neck obtained on one manufacturer's device to the expected value on another manufacturer's device (see Appendix II) (63). The margin of error in such conversions is still too great to use such values in following a patient over time, however. Such values should only be viewed as "ball park" figures. Another set of formulas makes possible the conversion of any manufacturer's BMD value at the lumbar spine or total hip to a second value called the "standardized bone mineral density" (sBMD see Appendix II) (63,64). The sBMD is always reported in mg/cm2 to distinguish it from the manufacturer's BMD, which is reported in g/cm2.

Perhaps the most significant advance seen with DXA compared to DPA is the marked improvement in precision. Expressed as a coefficient of variation, short-term precision in normal subjects has been reported as low as 0.9% for the PA lumbar spine and 1.4% for the femoral neck (56). Precision studies over the course of 1 year have reported values of 1% for the PA lumbar spine and 1.7 to 2.3% for the femoral neck (59).

Radiation exposure is extremely low for all types of DXA scans. Expressed as skin dose, radiation exposure during a PA lumbar spine or proximal femur study is only 2 to 5 mrem.11 The biologically important effective dose or whole-body equivalent dose is only 0.1 mrem (65).

DXA has been used in prospective studies to predict fracture risk. In one of the largest studies of its kind, DXA studies of the proximal femur were demonstrated to have the greatest short-term predictive ability for hip fracture compared to measurements at other sites with SPA or DPA (19).

DXA central devices are called "pencil-beam" or "fan-array" scanners. Examples of pencil-beam scanners are the Lunar DPX® Plus, DPX®-L, DPX-IQ™, DPX®-SF, DPX®-A, DPX-MD™, DPX-MD+™ and DPX-NT™, the Hologic QDR® 1000 and QDR® 2000 and the Norland XR-36™, XR-46™, Excell™ and Excell™plus.12 Examples of fan-array DXA scanners are the Lunar Expert® and Prodigy™ and the Hologic QDR® 4500 A, QDR® 4500 C, QDR® 4500 W, QDR® 4500 SL, Delphi™, and Discovery™. The difference between the pencil-beam and fan-array scanners is illustrated in Figs. 1-9 and 1-10. Pencil-beam scanners employ a collimated or narrowed X-ray beam (narrow like a pencil) that moves in tandem in a rectilinear pattern with the detector(s). Fan-array scanners utilize a much broader or fan-shaped beam and an array of detectors, so that an entire scan line can be instantly quantified. Scan times are reduced to as short as 10 seconds for a PA study of the lumbar spine. Image resolution is also enhanced with the fan-array scanners as shown in the extraordinary images in Fig. 1-11. This has created a new application for bone densitometry scanning called

10 See Chapter 5 for a detailed discussion of the difference in values obtained using central devices from different manufacturers, conversions equations, and the development of the sBMD.

11 See Chapter 13 for a listing of radiation dose according to device and scan type.

12 Specific descriptions and photographs of these scanners can be found in Chapter 13.

Fig. 1-9. Pencil-beam DXA densitometers. The single detector or sequential detectors move in tandem with the narrowed X-ray beam in a rectilinear scan path.
Fig. 1-10. Fan-array DXA densitometers. An array of detectors and fan-shaped beam make possible the simultaneous acquisition of data across an entire scan line.
Fig. 1-11. Images from the fan-array imaging densitometer, the Lunar EXPERT-XL. Images courtesy of GE Medical Systems, Madison, WI.
Fig. 1-12. LVA™ image acquired on the Lunar Prodigy™. A fracture is apparent at T12. A safety pin is also seen over the anterior chest. Case courtesy of GE Medical Systems, Madison, WI.

morphometric X-ray absorptiometry (MXA). With MXA, images of the spine obtained in the lateral projection can be used for computer analysis of the vertebral dimensions and diagnosis of vertebral fracture. Fan-array scanners have also been developed to image the lateral spine in its entirety to allow a visual assessment of vertebral size and shape. Examples of scanners with this capability are the Hologic Delphi™, Discovery™, and the Lunar Prodigy™. Figures 1-12 and 1-13 are lateral spine images from the Lunar Prodigy™. In the LVA™ (Lateral Vertebral Assessment)23 image in Fig. 112 a fracture is suggested at T12. In Fig. 1-13, the dimensions of the suspect vertebra are measured with morphometry. Figs. 1-14 and 1-15 are IVA™ (Instant Vertebral Assessment) images from the Hologic Delphi™. No fractures are apparent in Fig. 114. Note the multiple thoracic deformities in Fig. 1-15.

DXA has effectively replaced DPA in both research and clinical practice. The shortened scan times, improved image resolution, lower radiation dose, improved precision, application to more skeletal sites, and lower cost of operation with DXA have relegated DPA to an honored place in densitometry history.

13 This application on newer GE Medical Systems devices is now called DVA (Dual-energy Vertebral Assessment). An image of the spine in the PA projection can be obtained in addition to the lateral view with DVA.

Fig. 1-13. LVA™ image acquired on the Lunar Prodigy™. Morphometric software allows the user to define the vertebral edges and measure the vertebral heights to quantitatively diagnose fracture. Case courtesy of GE Medical Systems, Madison, WI.

Peripheral DXA

DXA technology is also employed in portable devices dedicated to the measurement of one or two appendicular sites. As such, these devices are characterized as "peripheral" (pDXA) devices. Because these devices employ dual-energy X ray, they do not require a water bath or tissue-equivalent gel surrounding the region of the skeleton being studied. As a consequence, they are somewhat easier to maintain and use than SXA devices. Examples of pDXA units are the Lunar PIXI®, the Norland pDEXA® and the Norland Apollo™, the Schick accuDEXA™, and the Osteometer DexaCare® DTX-200 and G4. These devices are discussed in detail in Chapter 13.

Single-Energy X-Ray Absorptiometry

SXA is the X-ray based counterpart of SPA, much as DXA is the X-ray-based counterpart of DPA. SXA units were used to measure BMD in the distal radius and ulna and calcaneus. Like their DXA counterparts, SXA units did not utilize radioactive isotopes but did require a water bath or tissue-equivalent gel surrounding the region of the skeleton being measured. The accuracy and precision of SXA were comparable to SPA (66). With the development of portable DXA devices for the measurement of forearm and heel bone density that do not require a water bath or tissue-equivalent gel, SXA is largely obsolete, just like its predecessor SPA.

Fig. 1-14. IVA™ image acquired on the Hologic Delphi™. No fractures are apparent in the thoracic and lumbar spine although aortic calcification is seen anterior to the lumbar spine. Case courtesy of Hologic Inc., Bedford, MA.

Fig. 1-15. IVA™ image acquired on the Hologic Delphi™. There are multiple deformities in the thoracic spine as well as osteophytes in the lower lumbar spine. Aortic calcification is also seen anterior to the lumbar spine. Case courtesy of Hologic Inc., Bedford, MA.

Fig. 1-14. IVA™ image acquired on the Hologic Delphi™. No fractures are apparent in the thoracic and lumbar spine although aortic calcification is seen anterior to the lumbar spine. Case courtesy of Hologic Inc., Bedford, MA.

Fig. 1-15. IVA™ image acquired on the Hologic Delphi™. There are multiple deformities in the thoracic spine as well as osteophytes in the lower lumbar spine. Aortic calcification is also seen anterior to the lumbar spine. Case courtesy of Hologic Inc., Bedford, MA.

Quantitative Computed Tomography

Although QCT is a photon absorptiometric technique like SPA, SXA, DPA, and DXA, it is unique in that it provides a three-dimensional or volumetric measurement of bone density and a spatial separation of trabecular from cortical bone. In 1976, Ruegsegger et al. (67) developed a dedicated peripheral QCT scanner using 125I for measurements of the radius. Cann and Genant (68,69) are credited with adapting commercially available CT scanners for the quantitative assessment of spinal bone density.

It is this approach that has received the most widespread use in the United States, although dedicated CT units for the measurement of the peripheral skeleton, or pQCT units, are in use in clinical centers. QCT studies of the spine utilize a reference standard or phantom that is scanned simultaneously with the patient. The phantom contains varying concentrations of K2HPO4 and is placed underneath the patient during the study. A scout view (see Fig. 1-16) is required for localization and then an 8- to 10-mm thick slice is measured through the center of two or more vertebral bodies that are generally selected from T12 to L3 (70). A region of interest within the anterior portion of the vertebral body is analyzed for bone density and is reported as mg/cm3 K2HPO4 equivalents (see Fig. 1-17). This region of interest is carefully placed to avoid the cortical shell of the vertebral body. The result is a three-dimensional trabecular density unlike the two-dimensional areal mixed cortical and trabecular densities reported with PA studies of the spine utilizing DPA or DXA.

A study of the spine with QCT requires about 30 minutes (35). The skin radiation dose is generally 100 to 300 mrem. This overestimates the biologically important effective dose because only a small portion of marrow is irradiated during a QCT study of the spine (65). The effective dose or whole-body equivalent dose is generally in the range of only 3 mrem (30 ^Sv). The localizer scan that precedes the actual QCT study will add an additional 3 mrem to the effective dose. These values are quite acceptable in the context of natural background radiation of approximately 20 mrem per month. Older CT units, that by their design are unable to utilize low kVp settings for QCT studies, may deliver doses 3 to 10 times higher.

The accuracy of QCT for measurements of spine BMD can be affected by the presence of marrow fat (70-72). Marrow fat increases with age, resulting in an increasingly large error in the accuracy of spine QCT measurements in older patients. The accuracy of QCT is reported to range from 5 to 15%, depending on the age of the patient and percentage of marrow fat. The presence of marrow fat results in an underestimation of bone density in the young of about 20 mg/cm3 and as much as 30 mg/cm3 in the elderly (70). The error introduced by marrow fat can be partially corrected by applying data on vertebral marrow fat with aging originally developed by Dunnill et al. (73). In an attempt to eliminate the error introduced by marrow fat, dual-energy QCT (DEQCT) was developed by Genant and Boyd (74). DEQCT clearly reduced the error introduced by the presence of marrow fat to as low as 1.4% in cadaveric studies (71,72). In vivo, the accuracy with DEQCT is 3 to 6% (35,70). Radiation dose with DEQCT is increased approximately 10-fold compared to regular or single-energy QCT (SEQCT) and precision is not as good. The precision of SEQCT for vertebral measurements in expert hands is 1 to 3% and for DEQCT, 3 to 5% (70,75).

The measurement of bone density in the proximal femur with QCT is not readily available. Using both dedicated QCT and standard CT units, investigators have attempted to utilize QCT for measurements of the proximal femur but this capability remains restricted to a few research centers (76,77).

QCT of the spine has been used in studies of prevalent osteoporotic fractures and it is clear that such measurements can distinguish osteoporotic individuals from normal individuals as well or even better than DPA (78-81). Fractures are rare with values above 110 mg/cm3 and extremely common below 60 mg/cm3 (82). Because QCT can isolate and measure trabecular bone, which is more metabolically active than cortical bone, rates of change in disease states observed with QCT spine measurements tend to

Fig. 1-16. QCT-5000™ scout image. Reproduced courtesy of Image Analysis Inc., Columbia, KY.
Fig. 1-17. QCT-5000™ axial spine image. This is a three-dimensional volumetric measurement, reported in mg/cm3 or mg/cc. The L2 BMD shown here is 120.2 mg/cc. This measurement is 100% trabecular. Reproduced courtesy of Image Analysis Inc., Columbia, KY.

be greater than those observed with PA spine studies performed with DPA or DXA (68,83). This greater magnitude of change partially offsets the effects of the poorer precision seen with QCT compared to DXA.24 The correlations between spine bone density measurements with QCT and skeletal sites measured with other techniques are statistically significant but too weak to allow accurate prediction of bone density at another site from measurement of the spine with QCT (26,80,81). This is no different however, from attempting to use BMD at the spine obtained with DXA to predict BMD at other skeletal sites.

Peripheral QCT

pQCT is becoming more widely available. pQCT devices are utilized primarily for the measurement of bone density in the forearm. Like QCT scans of the spine, pQCT makes possible true three-dimensional or volumetric measurements of bone density in the forearm, which may be particularly useful when the size of the bone is changing, as in pediatric populations. Information on a commercially available pQCT device, the Stratec XCT 2000™, can be found in Chapter 13.

quantitative ultrasound bone densitometry

Research in quantitative ultrasound (QUS) bone densitometry has been ongoing for more than 40 years. Only in the last few years, however, has QUS begun to play a role in the clinical evaluation of the patient. Ultrasound technologies in clinical medicine have traditionally been imaging technologies used, for example, to image the gall bladder or the ovaries. Like photon absorptiometric technologies, however, the application of ultrasound in bone densitometry is not primarily directed at producing an image of the bone. Instead, a quantitative assessment of bone density is desired with the image being secondary in importance.

In theory, the speed with which sound passes through bone is related not only to the density of the bone, but to the quality of the bone as well. Both bone density and bone quality determine a bone's resistance to fracture. Therefore, the speed of sound through bone can be related to the risk of fracture. These relationships can be illustrated mathematically. For example, the bone's ability to resist fracture (R) can be described as the amount the bone deforms when it is subjected to a force (F) that is moderated by the bone's ability to resist that force, the elastic modulus (E) as shown in Equation 1.

Studies have shown that the elastic modulus, E, is determined by both bone density and bone quality. Mathematically this is represented in Equation 2, where K is a constant representing bone quality and p represents bone density.

From such an equation, it becomes clear that the bone's ability to resist a force and not fracture is determined by changes in bone density and bone quality. When ultrasound

14 See Chapter 11 for a detailed discussion on the interaction between precision and rate of change in determining the time interval required between measurements to demonstrate significant change.

passes through a material, the velocity of the sound wave is also related to the elastic modulus (84,85) and density of the material as shown in Equation 3.

When Equations 2 and 3 are combined, it becomes clear that the velocity of ultrasound through bone is directly related to the square root of the product of bone density and bone quality.

The velocity with which ultrasound passes through normal bone is quite fast and varies depending on whether the bone is cortical or trabecular. Speeds of 3000 to 3600 m/s are typical in cortical bone with speeds of 1650 to 2300 m/s typical of trabecular bone.

In order to calculate velocity, ultrasound densitometers must measure the distance between two points and the time required for the sound wave to travel between these two points. The velocity is reported as the speed of sound (SOS). Higher values of SOS indicate greater values of bone density.

A second ultrasound parameter is broadband ultrasound attenuation (BUA). This parameter is reported in decibels per megahertz (dB/MHz). BUA is perhaps best understood using the analogy of a child's slinky toy. When the toy is stretched out and then suddenly released, the energy imparted to the rings by stretching them causes the rings to oscillate for a period of time, with the oscillations becoming progressively less and then finally stopping as the energy is lost. The same thing happens to the sound wave as it passes through bone. Some of the energy is lost from the sound wave and the oscillations of the sound wave are diminished. How much energy is lost is again related to the density of the bone and to architectural qualities such as porosity and trabecular connectivity (84,85). Like SOS, higher BUA values indicate greater bone density.

Most devices report both SOS and BUA. However, one manufacturer has mathematically combined SOS and BUA into a proprietary index called the Stiffness Index. Another manufacturer reports a proprietary index called the Quantitative Ultrasound Index (QUI) and an estimated BMD that is derived from the measurements of SOS and BUA. QUS devices are considered peripheral devices and are generally quite portable. They employ no ionizing radiation, unlike their SXA or DXA peripheral counterparts. The calcaneus is the most common skeletal site assessed with QUS, but devices exist that can be applied to the radius, finger, and tibia. In heel QUS measurements, heel width apparently has little if any effect on BUA but may have a slight effect on SOS (86). Most ultrasound devices require some type of coupling medium between the transducers and the bone. This is often accomplished with water when the heel is placed directly into a water bath. Ultrasound gel may be used in place of direct contact with water for heel measurements and measurements at other skeletal sites. Systems that utilize water baths into which the foot is placed are called "wet" systems. Systems that do not require water submersion but utilize gel instead, are called "dry" systems. There is one system for the heel in which neither water submersion or gel are required, making it truly dry. The Lunar Achilles+™, the Lunar Achilles Express™, the Lunar Insight™,

the Sunlight Omnisense™ 7000S, the Quidel QUS-2™, the McCue C.U.B.A. Clinical™, the Hologic Sahara™ Clinical Bone Sonometer, and the Osteometer Ultrasure DTU-1 are all examples of QUS densitometers currently available for clinical use. These devices are discussed in more detail in Chapter 13.

The technical differences between QUS devices from various manufacturers are even greater than those seen with DXA devices. Different frequency ranges and transducer sizes may be employed from device to device. Within the same skeletal site, slightly different regions of interest may be measured. As a consequence, values obtained on one QUS device are not necessarily comparable to values obtained on another QUS device.

The physics of ultrasound suggest that it should provide information about the bone that goes beyond a simple measurement of mass or density. Clinical research has tended to confirm this assumption, although perhaps not to the extent that was originally hoped. In a very large study of 5662 older women, both SOS and BUA predicted the risk of hip fracture as well or better than did measurements of BMD at the femoral neck using DXA (87). Similar findings were reported in the Study of Osteoporotic Fractures by Bauer et al. (88).

The precision of QUS measurements is generally excellent. In addition, because of the speed with which measurements can be made and the lack of any ionizing radiation, measurements can be made in duplicate or triplicate at any one examination. The average value of such replicate studies can be used, which dramatically improves precision. In a study from Njeh et al. (89) in which the precision of six different calcaneal QUS devices was determined, the short-term precision for SOS, expressed as the root-mean-square percent coefficient of variation (RMS-%CV) ranged from 0.11 to 0.42. For BUA, the RMS-%CV ranged from 1.39 to 6.30. Typically, better precision values are seen for SOS than for BUA.


1. Dennis J. A new system of measurement in X-ray work. Dental Cosmos 1897;39:445-454.

2. Price WA. The science of dental radiology. Dental Cosmos 1901;43:483-503.

3. Johnston CC, Epstein S. Clinical, biochemical, radiographic, epidemiologic, and economic features of osteoporosis. Orthop Clin North Am 1981;12:559-569.

4. Aitken M. Measurement of bone mass and turnover. Osteoporosis in clinical practice. Bristol: John Wright & Sons Ltd, 1984:19-20.

5. Singh J, Nagrath AR, Maini PS. Changes in trabecular pattern of the upper end of the femur as an index of osteoporosis. J Bone Joint Surg Am 1970;52-A:457-467.

6. Bohr H, Schadt O. Bone mineral content of femoral bone and lumbar spine measured in women with fracture of the femoral neck by dual photon absorptiometry. Clin Ortho 1983;179:240-245.

7. Nordin BEC. Osteoporosis with particular reference to the menopause. In: Avioli LV, ed. The osteoporotic syndrome. New York: Grune & Stratton, 1983:13-44.

8. Shimmins J, Anderson JB, Smith DA, et al. The accuracy and reproducibility of bone mineral measurements "in vivo." (a) The measurement of metacarpal mineralisation using an X-ray generator. Clin Radiol 1972;23:42-46.

9. Exton-Smith AN, Millard PH, Payne PR, Wheeler EF. Method for measuring quantify of bone. Lancet 1969;2:1153-1154.

10. Dequeker J. Precision of the radiogrammetric evaluation of bone mass at the metacarpal bones. In: Dequeker J, Johnston CC, eds. Non-invasive bone measurements: methodological problems. Oxford: IRL Press, 1982:27-32.

11. Aitken JM, Smith CB, Horton PW, et al. The interrelationships between bone mineral at different skeletal sites in male and female cadavera. J Bone Joint Surg Br 1974;56B:370-375.

12. Meema HE, Meindok H. Advantages of peripheral radiogrammetry over dual-photon absorptiometry of the spine in the assessment of prevalence of osteoporotic vertebral fractures in women. J Bone Miner Res 1992;7:897-903.

13. Bywaters EGL. The measurement of bone opacity. Clin Sci 1948;6:281-287.

14. Barnett E, Nordin BEC. Radiologic assessment of bone density. 1. The clinical and radiological problem of thin bones. Br J Radiol 1961;34:683-692.

15. Bouxsein ML, Palermo L, Yeung C, Black DM. Digital X-ray radiogrammetry predicts hip, wrist and vertebral fracture risk in elderly women: a prospective analysis from the Study of Osteoporotic Fractures. Osteoporos Int 2002;12:358-365.

16. Cummings S, Black D, Nevitt M, et al. Appendicular bone density and age predict hip fractures in women. JAMA 1990;263:665-668.

17. Mack PB, Brown WN, Trapp HD. The quantitative evaluation of bone density. Am J Roentgenol Rad Ther 1949;61:808-825.

18. Vose GP, Mack PB. Roentgenologic assessment of femoral neck density as related to fracturing. Am J Roentgenol Rad Ther Nucl Med. 1963;89:1296-1301.

19. Cummings SR, Black DM, Nevitt MC, et al. Bone density at various sites for prediction of hip fractures. Lancet 1993;341:72-75.

20. Mazess RB. Noninvasive methods for quantitating trabecular bone. In: Avioli LV, ed. The osteoporotic syndrome. New York: Grune & Stratton, 1983:85-114.

21. Mack PB, O'Brien AT, Smith JM, Bauman AW. A method for estimating degree of mineralization of bones from tracings of roentgenograms. Science 1939;89:467.

22. Mack PB, Vogt FB. (1971) Roentgenographic bone density changes in astronauts during representative Apollo space flight. Am J Roentgenol Rad Ther Nucl Med 1971;113:621-633.

23. Cosman F, Herrington B, Himmelstein S, Lindsay R. Radiographic absorptiometry: a simple method for determination of bone mass. Osteoporos Int 1991;2:34-38.

24. Yates AJ, Ross PD, Lydick E, Epstein RS. Radiographic absorptiometry in the diagnosis of osteoporosis. Am J Med 1995;98:41S-47S.

25. Yang S, Hagiwara S, Engelke K, et al. Radiographic absorptiometry for bone mineral measurement of the phalanges: precision and accuracy study. Radiology 1994;192:857-859.

26. Kleerekoper M, Nelson DA, Flynn MJ, Pawluszka AS, Jacobsen G, Peterson EL. Comparison of radiographic absorptiometry with dual-energy X-ray absorptiometry and quantitative computed tomography in normal older white and black women. J Bone Miner Res 1994;9:1745-1749.

27. Mussolino ME, Looker AC, Madans JH, et al. Phalangeal bone density and hip fracture risk. Arch Intern Med 1997;157:433-438.

28. Huang C, Ross PD, Yates AJ, et al. Prediction of fracture risk by radiographic absorptiometry and quantitative ultrasound: a prospective study. Calcif Tissue Int 1998;6:380-384.

29. Cameron JR and Sorenson G. Measurements of bone mineral in vivo: an improved method. Science 1963;142:230-232.

30. Vogel JM. Application principles and technical considerations in SPA. In: Genant HK, ed. Osteoporosis update 1987. San Francisco: University of California Printing Services, 1987:219-231.

31. Johnston CC. Noninvasive methods for quantitating appendicular bone mass. In: Avioli L, ed. The osteoporotic syndrome. New York: Grune & Stratton, 1983:73-84.

32. Barden HS, Mazess RB. Bone densitometry of the appendicular and axial skeleton. Top Geriatric Rehabi. 1989;4:1-12.

33. Kimmel PL. Radiologic methods to evaluate bone mineral content. Ann Intern Med 1984;100:908-911.

34. Steiger P, Genant HK. The current implementation of single-photon absorptiometry in commercially available instruments. In: Genant HK, ed. Osteoporosis update 1987. San Francisco: University of California Printing Services, 1987:233-240.

35. Chesnut CH. Noninvasive methods for bone mass measurement. In: Avioli L, ed. The Osteoporotic syndrome. 3rd ed. New York: Wiley-Liss, 1993:77-87.

36. Gardsell P, Johnell, O, Nilsson BE. The predictive value of bone loss for fragility fractures in women: a longitudinal study over 15 years. Calcif Tissue Int 1991;49:90-94.

37. Hui SL, Slemenda CW, Johnston CC. Baseline measurement of bone mass predicts fracture in white women. Ann Intern Med 1989;111:355-361.

38. Ross PD, Davis JW, Vogel JM, Wasnich RD. A critical review of bone mass and the risk of fractures in osteoporosis. Calcif Tissue Int 1990;46:149-161.

39. Melton LJ, Atkinson EJ, O'Fallon WM, Wahner HW, Riggs BL. Long-term fracture prediction by bone mineral assessed at different skeletal sites. J Bone Miner Res 1993;8:1227-1233.

40. Black DM, Cummings SR, Genant HK, Nevitt MC, Palermo L, Browner W. Axial and appendicular bone density predict fracture in older women. J Bone Miner Res 1992;7:633-638.

41. Nord, R.H. Technical considerations in DPA. In: Genant HK., ed. Osteoporosis update 1987. San Francisco: University of California Printing Services, 1987:203-212.

42. Dunn WL, Wahner HW, Riggs BL. Measurement of bone mineral content in human vertebrae and hip by dual photon absorptiometry. Radiology 1980;136:485-487.

43. Reed GW. The assessment of bone mineralization from the relative transmission of 241Am and 137Cs radiations. Phys Med Biol 1966;11:174.

44. Roos B, Skoldborn H. Dual photon absorptiometry in lumbar vertebrae. I. Theory and method. Acta Radiol Ther Phys Biol 1974;13:266-290.

45. Mazess RB, Ort M, Judy P. Absorptiometric bone mineral determination using 153Gd. In: Cameron JR, ed. Proceedings of bone measurements conference. U.S. Atomic Energy Commission, 1970:308-312.

46. Wilson CR, Madsen M. Dichromatic absorptiometry of vertebral bone mineral content. Invest Radiol 1977;12:180-184.

47. Madsen M, Peppler W, Mazess RB. Vertebral and total body bone mineral content by dual photon absorptiometry. Calcif Tissue Res 1976;2:361-364.

48. Wahner WH, Dunn WL, Mazess RB, et al. Dual-photon Gd-153 absorptiometry of bone. Radiology 1985;156:203-206.

49. Lindsay R, Fey C, Haboubi A. Dual photon absorptiometric measurements of bone mineral density increase with source life. Calcif Tissue Int 1987;41:293-294.

50. Cummings SR, Black DB. Should perimenopausal women be screened for osteoporosis? Ann Intern Med 1986;104:817-823.

51. Drinka PJ, DeSmet AA, Bauwens SF, Rogot A. The effect of overlying calcification on lumbar bone densitometry. Calcif Tissue Int 1992;50:507-510.

52. Curry TS, Dowdey JE, Murry RC. Christensen's physics of diagnostic radiology. Philadelphia: Lea & Febiger, 1990:1-522.

53. Rupich RC, Griffin MG, Pacifici R, Avioli LV, Susman N. Lateral dual-energy radiography: artifact error from rib and pelvic bone. J Bone Miner Res 1992;7:97-101.

54. Louis O, Van Den Winkel P, Covens P, Schoutens A, Osteaux M. Dual-energy X-ray absorptiometry of lumbar vertebrae: relative contribution of body and posterior elements and accuracy in relation with neutron activation analysis. Bone 1992;13:317-320.

55. Peel NFA, Johnson A, Barrington NA, Smith TWD, Eastell R. Impact of anomalous vertebral segmentation of measurements of bone mineral density. J Bone Miner Res 1993;8:719-723.

56. Lees B, Stevenson JC. An evaluation of dual-energy X-ray absorptiometry and comparison with dualphoton absorptiometry. Osteoporos Int. 1992;2:146-152.

57. Kelly TL, Slovik DM, Schoenfeld DA, Neer RM. Quantitative digital radiography versus dual photon absorptiometry of the lumbar spine. J Clin Endocrinol Metab 1988;76:839-844.

58. Holbrook TL, Barrett-Connor E, Klauber M, Sartoris D. A population-based comparison of quantitative dual-energy X-ray absorptiometry with dual-photon absorptiometry of the spine and hip. Calcif Tissue Int 1991;49:305-307.

59. Pouilles JM, Tremollieres F, Todorovsky N, Ribot C. Precision and sensitivity of dual-energy X-ray absorptiometry in spinal osteoporosis. J Bone Miner Res 1991;6:997-1002.

60. Laskey MA, Cirsp AJ, Cole TJ, Compston JE. Comparison of the effect of different reference data on Lunar DPX and Hologic QDR-1000 dual-energy X-ray absorptiometers. Br J Radiol 1992;65: 1124-1129.

61. Pocock NA, Sambrook PN, Nguyen T, Kelly P, Freund J, Eisman J. Assessment of spinal and femoral bone density by dual X-ray absorptiometry: comparison of Lunar and Hologic instruments. J Bone Miner Res 1992;7:1081-1084.

62. Lai KC, Goodsitt MM, Murano R, Chesnut CC. A comparison of two dual-energy X-ray absorptiometry systems for spinal bone mineral measurement. Calcif Tissue Int 1992;50:203-208.

63. Genant HK, Grampp S, Gluer CC, et al. Universal standardization for dual X-ray absorptiometry: patient and phantom cross-calibration results. J Bone Miner Res 1994;9:1503-1514.

64. Hanson J. Standardization of femur BMD. J Bone Miner Res 1997;12:1316-1317.

65. Kalender WA. Effective dose values in bone mineral measurements by photon-absorptiometry and computed tomography. Osteoporos Int 1992;2:82-87.

66. Kelly TL, Crane G, Baran DT. Single x-ray absorptiometry of the forearm: precision, correlation, and reference data. Calcif Tissue Int 1994;54:212-218.

67. Ruegsegger P, Elsasser U, Anliker M, Gnehn H, Kind H, Prader A. Quantification of bone mineralisation using computed tomography. Radiology 1976;121:93-97.

68. Genant HK, Cann CE, Ettinger B, Gorday GS. Quantitative computed tomography of vertebral spon-giosa: a sensitive method for detecting early bone loss after oophorectomy. Ann Intern Med 1982;97: 699-705.

69. Cann CE, Genant HK. Precise measurement of vertebral mineral content using computed tomography. J Comput Assist Tomogr 1980;4:493-500.

70. Genant HK, Block JE, Steiger P, Gluer C. Quantitative computed tomography in the assessment of osteoporosis. In: Genant HK, ed. Osteoporosis update 1987. San Francisco: University of California Printing Services, 1987:49-72.

71. Laval-Jeantet AM, Roger B, Bouysse S, Bergot C, Mazess RB. Influence of vertebral fat content on quantitative CT density. Radiology 1986;159:463-466.

72. Reinbold W, Adler CP, Kalender WA, Lente R. Accuracy of vertebral mineral determination by dual-energy quantitative computed tomography. Skeletal Radiol 1991;20:25-29.

73. Dunnill MS, Anderson JA, Whitehead R. Quantitative histological studies on age changes in bone. J Pathol Bacteriol 1967;94:274-291.

74. Genant HK, Boyd D, Quantitative bone mineral analysis using dual energy computed tomography. Invest Radiol 1977;12:545-551.

75. Cann CE. Quantitative computed tomography for bone mineral analysis: technical considerations. In: Genant HK, ed. Osteoporosis update 1987. San Francisco: University of California Printing Services, 1987:131-144.

76. Sartoris DJ, Andre M, Resnick C, Resnick D. Trabecular bone density in the proximal femur: quantitative CT assessment. Radiology 1986;160:707-712.

77. Reiser UJ, Genant HK. Determination of bone mineral content in the femoral neck by quantitative computed tomography. 70th Scientific Assembly and Annual Meeting of the Radiological Society of North America, Washington, DC, 1984.

78. Gallagher C, Golgar D, Mahoney P, McGill J. Measurement of spine density in normal and osteoporotic subjects using computed tomography: relationship of spine density to fracture threshold and fracture index. J Comput Assist Tomogr 1985;9:634-635.

79. Raymaker JA, Hoekstra O, Van Putten J, Kerkhoff H, Duursma SA. Osteoporosis fracture prevalence and bone mineral mass measured with CT and DPA. Skeletal Radiol 1986;15:191-197.

80. Reinbold WD, Reiser UJ, Harris ST, Ettinger B, Genant HK. Measurement of bone mineral content in early postmenopausal and postmenopausal osteoporotic women. A comparison of methods. Radiology 1986;160:469-478.

81. Sambrook PN, Bartlett C, Evans R, Hesp R, Katz D, Reeve J. Measurement of lumbar spine bone mineral: a comparison of dual photon absorptiometry and computed tomography. Br J Radiol 1985;58: 621-624.

82. Genant HK, Ettinger B, Harris ST, Block JE, Steiger P. Quantitative computed tomography in assessment of osteoporosis. In: Riggs BL, Melton LJ, eds. Osteoporosis: etiology, diagnosis and management. New York: Raven Press, 1988:221-249.

83. Richardson ML, Genant HK, Cann CE, et al. Assessment of metabolic bone disease by quantitative computed tomography. Clin Orth Rel Res 1985;195:224-238.

84. Gluer C, Wu C, Jergas M, Goldstein S, Genant H. Three quantitative ultrasound parameters reflect bone structure. Calcif Tissue Int 1994;55:46-52.

85. Nicholson P, Haddaway M, Davie M. The dependence of ultrasonic properties on orientation in human vertebral bone. Phys Med Biol 1994;39:1013-1024.

86. Njeh CF, Boivin CM, Langton CM. The role of ultrasound in the assessment of osteoporosis: a review. Osteoporos Int 1997;7:7-22.

87. Hans D, Dargent-Molina P, Schott AM, et al. Ultrasonographic heel measurements to predict hip fracture in elderly women: the EPIDOS prospective study. Lancet 1996;348:511-514.

88. Bauer DC, Gluer CC, Cauley JA, et al. Bone ultrasound predicts fractures strongly and independently of densitometry in older women: a prospective study. Arch Intern Med 1997;157:629-634.

89. Njeh CF, Hans D, Li J, et al. Comparison of six calcaneal quantitative ultrasound devices: precision and hip fracture discrimination. Osteoporos Int 2000;11:1051-1062.

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