Bone densitometer

In an x-ray bone densitometer, special calibration techniques are employed to accommodate variations. In one aspect, a bone-like calibration material is interposed and the system determines the calibration data from rays passing only through flesh. In another aspect, a rotating device carries the calibration material through the beam. The specific densitometer shown uses an x-ray tube operated at two different voltages to generate a pencil beam, the energy levels of the x-ray photons being a function of the voltage applied. An integrating detector is timed to integrate the detected signal of the patient-attenuated beam over each pulse, the signals are converted to digital values and a digital computer converts the set of values produced by the raster scan into a representation of the bone density of the patient. Multiple reference detectors with differing absorbers are used by the system to continuously correct for variation in voltage and current of the x-ray tube. Calibration is accomplished by the digital computer on the basis of passing the pencil beam through known bone-representing substance as the densitometer scans portions of the patient having bone and adjacent portions having only flesh. A set of detected signals affected by the calibration substance in regions having only flesh is compared by the computer with a set of detected signals unaffected by the calibration material.

The invention is an x-ray densitometer suitable to measure bone density, or 
density of bone like materials, in the human body, particularly in the 
spine and hip. Such measurements are useful, e.g. for determining whether 
patients are affected by osteoporosis. The invention uses a measurement 
technique which is an improvement over a related technique called 
dual-photon absorptiometry. Dual photon absorptiometry is based on the use 
of radioisotopic sources to provide photons of two different energies 
whereas the present invention uses an x-ray tube switched between two 
different voltages in order to generate a collimated beam of two different 
energies. 
An x-ray source is capable of producing an intensity of radiation about 
1000 times greater than conventional radioisotopic sources used for bone 
density measurements. If an x-ray source were successfully incorporated in 
a bone densitometer, an improvement in measurement time, resolution, 
accuracy, precision, and minimization of radiation dose might be effected. 
Prior efforts to use x-ray sources for bone densitometers have not been 
altogether successful. A major purpose of the present invention is to 
provide a successful bone densitometer and to achieve improved performance 
in all of the important categories by taking advantage of the high 
radiation intensity produced by an x-ray source. 
The invention achieves this objective by use of an x-ray tube which is 
moved in a 2-dimensional raster scanning pattern with a fixed relationship 
between the tube and a collimator and detector which move with it, with 
alternating high and low voltage levels being applied to the x-ray tube. 
In order to take optimal advantage of the high photon intensity provided by 
x-ray sources, the invention overcomes certain problems associated with 
using x-ray tubes for bone densitometry. Although x-ray sources are more 
intense than radioisotopic sources, they are also less stable because they 
vary (drift) with changes in the voltage and current supplied to them. In 
addition, x-ray tubes produce photons that have a broad range of energies 
whereas radioisotopic sources typically produce photons with only a few 
energies. 
These and other problems are met by a system which employs two reference 
detectors instead of one, means for providing frequent bone calibration, 
e.g. on every scan line, preferably on every point, and use of an 
integrating detector in the raster scan. 
An important feature of the invention is a calibration technique which 
determines the location of bone and then calibrates the system on the 
basis of x-ray data produced from non bone areas that lie close to the 
location of bone. Such calibration not only accomodates drift of the x-ray 
source but also other variations encountered, such as variation in body 
thickness from patient to patient. 
To summarize, according to one aspect of the invention, a bone densitometer 
is provided for measuring density of bone-like material in a patient who 
is held in fixed position, comprising an x-ray tube means having a power 
supply, detector means arranged on the opposite side of the patient to 
detect x-rays attenuated by the patient, means to effectively expose 
portions of the patient having bone and adjacent portions having only 
flesh (non-bone body substance), means for causing the beam to pass 
through a bone like calibration material in the course of the exposure, 
and signal processing means responsive to the output of the detector means 
to provide a representation of bone density of the patient (e.g., an x-ray 
film like picture of the patient, showing bone density distribution or 
calculated values representing bone density of the patient), the signal 
processing means adapted to respond to data based upon x-rays attenuated 
by the calibration material in regions having only flesh to calibrate the 
output of the detector means, thereby enabling the accommodation of drift 
in the x-ray tube, differences in patient thickness and other system 
variations. 
According to another aspect of the invention, a bone densitometer, e.g., 
having the features described above, further comprises a pencil beam 
collimator arranged to form and direct a pencil beam of x-rays through the 
patient, and the detector means, on the opposite side of the patient, is 
aligned with the collimator, the x-ray tubes, pencil beam collimator and 
detector means adapted to be driven in unison in an X Y raster scan 
pattern relative to the patient, and rotating means cause the beam to pass 
through bone-like calibration means. According to this aspect of the 
invention, the signal processing means is not limited to responding to 
data based upon x-rays attenuated by the calibration material in regions 
having only flesh to calibrate the output of the detector means. 
According to another aspect of the invention, a bone densitometer 
incorporates features of both of the above described aspects of this 
invention. 
In preferred embodiments of these aspects of the invention, the power 
supply of the bone densitometer is adapted to apply alternate high and low 
voltage levels to the x-ray tube; control means for the frequency of the 
voltage is related to the speed at which the x-ray tube, collimator and 
detector means are driven in scan motion and the beam width produced by 
the collimator to apply alternating high and low voltage levels to the 
x-ray tube at a frequency sufficiently high that at least one pair of high 
and low level exposures occurs during the short time period during which 
the pencil beam traverses a distance equal to about one beam width, 
preferably the bone densitometer being adapted to produce pairs of high 
and low voltage pulses at a rate of the order of sixty per second, the 
x-ray tube, collimator and detector means being driven along the scan at a 
rate of the order of one inch per second and the collimator produces a 
pencil beam of between about one and three millimeters in diameter; the 
x-ray beam passes through the bone-like calibration material at least once 
per scan line of a scan pattern for a period equal to at least the time 
during which one pixel of resolution is traversed, preferably the x-ray 
beam passes through the bone-like calibration material for the duration of 
every other high and low voltage pulse pair; the detector means comprises 
an integrating detector controlled to integrate the detected signal 
repeatedly over short time periods relative-to the time required to 
advance the x-ray scan pattern by one pixel of resolution, preferably an 
analog to digital converter being provided to convert each integrated 
value, to a digital signal and a digital computer means is provided for 
producing the representation of bone density cf the patient by processing 
the stream of the digital signals; and a reference system is provided 
having at least two reference detectors each provided with a different 
absorber, the reference system adapted to correct for both x-ray tube 
current and voltage changes, preferably the system adapted to correct the 
detected signal substantially on the basis of a function of the signals 
produced by the reference detectors and, where there are two of the 
reference detectors, the function being substantially a straight line 
defined by the detected signals of the reference detectors. 
The present invention makes it possible to perform bone density 
measurements more rapidly and with better resolution and accuracy than 
prior devices. Because it does not use radioisotopic sources, the user 
does not need to handle and replace radioactive materials which are 
dangerous and are strictly controlled by federal licensing regulations.

DESCRIPTION OF THE PREFERRED EMBODIMENT 
FIG. 1 shows the basic components of the x-ray densitometer. X-ray tube 1 
carried on x-y table arrangement 2 is energized by power supply 11 which 
is designed to alternate its voltage output rapidly between two levels 
called the "High Energy Level" (HEL) and the "Low Energy Level" (LEL). The 
HEL is typically 150 kilovolts and the LEL is typically 75 kilovolts. The 
x-rays emitted by the x-ray tube are collimated to form a pencil beam B by 
collimator 3. The pencil beam passes through a calibration disc 12 which 
rotates at a rate which is synchronized with the rate at which the power 
supply 11 switches between the HEL and LEL. The role played by the 
calibration disc will be described below. 
The x-ray pencil beam passes through, e.g., a female adult patient 4 under 
examination for possible presence of osteoporosis and impinges on a main 
radiation detector 5. Two reference detectors 7 and 8 which are similar in 
design to the main detector 5 are also shown in FIG. 1. The reference 
detectors 7 and 8 monitor the flux emitted by x-ray tube 1 and provide 
information used to correct the signal measured by main detector 5 for 
variations in the x-ray tube current and voltage. 
The reference detectors 7 and 8 measure radiation from the x-ray tube 1 
after it has traversed one of two x-ray absorbers 9 and 10. The two 
absorbers 9 and 10 are of substantially different thicknesses which are 
typically chosen to be representative of the x-ray attenuation of a thin 
patient and a heavy patient respectively. By using two reference detectors 
with different absorbers it is possible to monitor changes simultaneously 
in both the x-ray tube current and voltage. The reference detector 
measurements are used to correct the measurements made with the main 
detector 5 in order to compensate for these changes in tube current and 
voltage. The manner in which these corrections are made is described 
below. 
X ray tube 1, collimator 3, reference detectors 7 and 8 and absorbers 9 and 
10, calibration disc 11, and main detector 5 are all mechanically 
continuously scanned in the X direction across the body during which time 
signals from the main detector 5 and reference detectors 7 and 8 are 
digitized and stored in computer system 13. After each scan from right to 
left or left to right in FIG. 1 the assembly briefly stops moving in the X 
direction, and is indexed a small amount in the Y direction, out of the 
plane of FIG. 1. As a result of these motions, the pencil x-ray beam B 
undergoes a rectangular scanning pattern 18 such as shown in FIGS. 2 and 
2a. A modified rectangular scanning pattern shown as parallelogram pattern 
19 in FIG. 2b might also be used to measure a bone such as the neck or the 
femur, which is set at an angle in the human body. 
FIG. 1 illustrates the fixed relationship between the x-ray source 1 and 
main detector 5 during the scanning period throughout which patient 4 lies 
stationary on patient table 20. X ray tube 1, collimator 3, reference 
detectors 7, 8 and absorbers 9 and 10, and calibration disc 12 are mounted 
together in a single assembly called the source assembly 22. This assembly 
is mounted in turn below the patient on a conventional X-Y table 2. 
Separate stepping motors and lead screws are used to move the X Y table in 
the X direction and Y-direction respectively. The stepping motors and lead 
screws are of a type well known in the art and are not shown. 
The main detector 5 is mounted above the patient and in the preferred 
embodiment shown is rigidly attached by means of C arm 21 to the source 
assembly 22 so that x-ray pencil beam B and main detector 5 have a fixed 
relationship throughout the scan. The main detector 5, in alternative 
embodiments, could be driven with its own drive system in either the 
X-direction, Y-direction or both so long as it maintains the same fixed 
relationship to pencil beam B. 
In FIG. 2, a representation of the patient's spine 6 and adjacent portions 
of the body is shown. In general, pencil beam B scans from side to side 
across the patient's spine, and through flesh on either side of the spine, 
but does not pass beyond the outer dimensions of the adult patient 4. The 
total distance scanned from side to side (i.e. in the X-direction) is 
typically 5 inches and the total distance scanned from head to toe (i.e. 
in the Y-direction) is typically 5 inches. 
During the scanning period, the signals from detector 5 and from reference 
detectors 7 and 8 are digitized and stored in computer system 13. It is 
possible to calculate the bone density at each point in the scan pattern 
from these data using a method described in more detail below. Both the 
raw data and the calculated bone density can be displayed as an image 
using any one of a number of devices well known in the art. Such an image 
will resemble a conventional x-ray image or radiograph. In a preferred 
embodiment, the computer system 13 contains one such device known as a 
display processor which displays the image acquired in this manner on a 
television screen. Devices such as a display processor or other computer 
peripherals such as laser printers are well known devices for displaying 
images from digital data. 
FIG. 3 is an electronic block diagram of the bone densitometer showing the 
relationship between the different components of the system. X-ray photons 
striking the crystals in the scintillation detectors 5, 7, 8 generate 
optical radiation which is converted by the detector photomultiplier tubes 
into electrical currents. These in turn are amplified and converted to 
voltage levels by individual amplifiers 30. The amplifier outputs are 
integrated by respective integrators 32 for time periods that are 
controlled by the system timing control 36 about which more will be said. 
The output of the three integrators are digitized by an analog-to-digitial 
converter 34 and stored for processing in a small computer 13 such as an 
IBM PC or AT computer system. 
The system timing control 36 synchronizes the x-ray power supply pulsing 
and the signal integrators. For example, just before a HEL voltage is 
applied to the x-ray tube, all three integrators are reset to zero. As the 
HEL is applied to the x-ray tube, radiation is emitted and all three 
integrators begin to integrate signals. A short time later (typically 
1/120 second), the system timing control terminates the HEL voltage level, 
terminating the emission of x-radiation. This is immediately followed by a 
signal generated by the timing control which terminates the integration of 
detector signals and causes the A/D converter to digitize the integrator 
output and transfer the digital value to the computer system. A similar 
sequence of timing signals is then generated for the next LEL pulse after 
which the cycle is repeated. 
The system timing control 36, in addition to the functions described above, 
also provides a means to synchronize, via synchronizing circuit 37, the 
calibration disc motor 12a to assure that the rotation frequency of the 
calibration disc 12 and the x-ray pulsing frequency are locked together so 
that the timing relationships illustrated in FIG. 5 are maintained. The 
timing control in a preferred embodiment also provides a pulse sequence to 
the stepping motor controller 38 which drives X and Y-direction motors 2a 
and 2b and assures that every scan line in the x-ray image has exactly the 
same number and phasing of x-ray pulses. 
The computer 13 provides scan distance instructions to the stepping motor 
controller 38 and scan initiation instructions to the system timing 
control 36 and allows the operator to initiate, manipulate, and terminate 
the raster scan motion and x-ray generation by means of a standard 
keyboard. It records the digitized detector information, calculates bone 
density for each point in the raster scan pattern, and displays the 
resulting image using a standard computer display processor and television 
monitor. Hardcopy versions of the calculated and displayed bone density 
can be obtained with a standard printer interfaced to the computer. 
BEAM SIZE AND SWITCHING FREQUENCY 
During the scanning of the x-ray pencil beam B, the voltage on the x-ray 
tube 1 is switched between the HEL and LEL. A typical speed used to 
continuously scan across the patient from side to side is 1 inch per 
second and a typical switching frequency for the x-ray tube power supply 
is 60 cycles per second. In this case there will be 60 HEL pulses 
alternating with 60 LEL pulses generated during each one second of 
scanning. The signals from detectors 5, 7 and 8 are recorded separately 
for the HEL pulses and for the LEL pulses. For a scan speed of 1 inch per 
second, one HEL/LEL pair of measurements is made for every 1/60 of an inch 
(0.016 inch) traversed by the pencil beam. 
The cross sectional area of the pencil beam B is determined by the opening 
in collimator 3 and is typically 1-3 mm (0.040-0.120 inch) so that there 
are typically 21/2 to 8 HEL/LEL pulse pairs per beam width. One of the 
important features of the present invention is that there is at least 
about one pulse pair per beam width. As a result, the small region of the 
body sampled by the pencil beam during the HEL measurement will be 
essentially the same as the small region of the body sampled by the LEL 
measurement made 1/120 second later. 
USE OF INTEGRATING DETECTORS 
The detectors used in the preferred x-ray densitometer are of a type 
generally known as scintillation detectors, although use of a number of 
other types of detectors is also possible. A scintillation detector 
consists of a crystal material coupled to a photomultiplier tube. The 
crystal serves to convert x-ray radiation to optical radiation and the 
photomultiplier tube converts the optical radiation to an electronic 
signal. Solid state photodiodes coupled to x-ray fluorescent screens, 
ionization chambers, and other devices might also serve as radiation 
detectors for the present invention. 
In dual-photon bone densitometers using radioisotopes, scintillation 
detectors are also used as radiation detectors. However, in these devices, 
the detectors must detect individual x-ray photons and sort these photons 
into two separate channels corresponding to high-energy photons and 
low-energy photons. This requirement for performing a spectrum analysis on 
individually detected photons is dictated by the fact that the isotopic 
source emits both high-energy and low energy photons simultaneously. 
In the present invention, the scintillation detector used need not perform 
a spectrum analysis task by sorting photons into high and low-energy 
channels because the high-energy and low-energy photons are not emitted 
simultaneously. Rather the HEL and LEL voltages are generated alternating 
in time. The high-energy and low-energy photons are integrated separately 
over the duration of the HEL and LEL pulses respectively and are therefore 
recorded at different times. 
The ability to use energy integrating detectors rather than photon counting 
detectors is an important feature of the x-ray densitometer because it 
makes it possible to complete a patient scan in a short time. In order to 
measure bone density to a given accuracy, it is necessary to detect a 
resulting minimum number of photons because the statistical accuracy of a 
measurement is related, as is well known, to the square root of the number 
of detected photons. For example, at least 50-100 million photons are 
typically detected in a bone density measurement of the spine. 
The x-ray densitometer of the present invention completes a measurement 
scan in as little as 2 to 5 minutes. In order to record as many as 100 
million photons in 2 minutes, the detector must record on the order of 1 
million photons per second. By using energy integrating detectors rather 
than pulse counting detectors, the x-ray densitometer can record photon 
fluxes of 1 million per second or higher. (An energy integrating detector 
can easily record photon fluxes as high as 100 million photons per 
second.) The use of alternating high and low voltage pulsing of the x-ray 
tube, coupled with integrating detectors to measure the high-energy and 
low-energy signals produced, is an important feature of the present 
invention because it makes short scan times possible, which implies a 
shorter visit by patients and better use of capital equipment. 
THE REFERENCE DETECTORS AND ABSORBERS 
By use of two reference detectors, with different absorbers, small changes 
in both the x-ray tube current and applied voltage are effectively 
monitored along with the signals from the main detector. Just as the main 
detector integrates photons and supplies separate values for the HEL pulse 
and the LEL pulse, the reference detectors also supply separate reference 
values for each HEL pulse and LEL pulse. Each HEL or LEL measurement 
recorded by the main detector is corrected according to the following 
method. 
Let P1 and P2 be the percentage changes in output signal (from a HEL or LEL 
pulse) measured by the first and second reference detectors respectively 
due to a small variation in x-ray tube current and voltage. Let T1 and T2 
be the thicknesses of body tissue that attenuate the HEL or LEL pulses the 
same amount as do the first and second absorbers shown in FIG. 1 
respectively. Each main detector signal corresponds to a body thickness 
T0, and is corrected by a percentage P0 which is given by the formula 
(P0-P1)=(P2-P1)/(T2-T1)*(T0-T1). (This can be recognized as the formula 
for a straight line fit between the measured values of P1 and P2.) This 
correction compensates the main detector measurement for changes in both 
x-ray tube current and voltage. This feature enables the main detector 
signal to be corrected for fluctuations in x-ray tube current and voltage 
that would otherwise degrade the accuracy of the bone density calculation, 
unless the x-ray power supply is quite stable over the duration of one 
patient scan. In the case that the power supply is sufficiently stable, 
the use of the reference detectors may be omitted. Longer term variations 
in x-ray tube current and voltage are compensated by the use of the 
calibration disc now to be described. 
THE CALIBRATION DISC 
FIG. 4 is a plan view of the calibration disc 12 which is mounted such that 
the region of the disc near the circumference interrupts pencil beam B as 
the disc rotates. The calibration disc is synchronized to the switching 
frequency of the high voltage power supply. In a preferred embodiment, the 
power supply produces HEL and LEL pulses which are derived from the main 
power line frequency of 60 Hertz. The HEL and LEL pulses generated by the 
power supply in this embodiment are shown in FIG. 5. One pair of HEL and 
LEL pulses are generated every 1/60 of a second. 
In the preferred embodiment, the calibration disc is driven with a 
synchronous motor which rotates at a rate of exactly 30 revolutions per 
second and which is adjusted in phase such that four pre-defined quadrants 
of the disc (labeled Q1, Q2, Q3, and Q4 in FIG. 4) correspond to the HEL 
and LEL levels being generated by the power supply. More specifically, 
when Quadrant 1 is obstructing the pencil beam the voltage level is HEL, 
when Quadrant 2 is obstructing the pencil beam the voltage level is LEL, 
when Quadrant 3 is obstructing the beam the voltage level is again HEL, 
and finally when Quadrant 4 is obstructing the beam the voltage level is 
again LEL. The desired synchronization between the quadrants of the 
calibration disc and the HEL/LEL voltage pulses is illustrated in FIG. 5. 
In this preferred embodiment, the circumference of Quadrants 1 and 2 
consists of a material which has the same x-ray attenuation 
characteristics as bone. Both quadrants contain exactly the same amount of 
the bone-like calibration material 15 which typically amounts to about 1 
gram per square centimeter of material. As a result, every other HEL/LEL 
pulse pair recorded by the main detector is attenuated by a constant 
thickness of calibration bone, as the calibration bone rotates in and out 
of the x-ray beam. Using the calibration disc in this manner, four 
distinct types of measurements are periodically recorded from the main 
detector. These types and their abbreviations are (1) HEL and no 
calibration bone (H), (2) LEL and no calibration bone (L); (3) HEL with 
calibration bone (HB), and (4) LEL with calibration bone (LB). The four 
groups of measurements H, L, HB, and LB conditions are illustrated in FIG. 
5. 
One additional function of the calibration disc is worth noting. It is 
possible to make use of the same disc for the additional purpose of 
providing different x-ray filtration for the HEL and LEL x-ray beams. For 
example, in a preferred embodiment, the LEL beam is left unfiltered 
whereas the HEL beam is filtered with 1 mm of copper. The purpose of the 
copper filtration is to attenuate the HEL beam, which is typically of 
considerably higher intensity than the LEL beam because it suffers less 
attenuation in tissue. By attenuating the HEL beam, it is possible to 
avoid unnecessary x-ray exposure to the patient and thereby lower the dose 
without substantially affecting the accuracy of the final bone 
measurement. 
In FIG. 4, in a preferred embodiment, Quadrants 1 and 3 contain the copper 
filtration in the form of a constant thickness sheet of copper 20 for the 
HEL beam and Quadrants 2 and 4 contain no filtration (or possibly another 
filtration material) for the LEL beam. The use of a rotating wheel to 
provide different x-ray filtrations is another advantage of the present 
invention, although the main purpose of the calibration disc is to provide 
continuous calibration of the densitometer as explained below. 
CALCULATION OF BONE DENSITY 
The method used to calculate bone density with high accuracy is based on 
dual photon absorptiometry calculations as described in prior 
publications. (See for example: "Noninvasive Bone Mineral Measurements," 
by Heinz W. Wahner, William L. Dunn, and B. Lawrence Riggs, Seminars in 
Nuclear Medicine, Vol. XIII, No. 3, 1983.) The x-ray densitometer 
described here uses a modification of the established method which makes 
use of the calibration disc bone-like material to obtain an absolute 
reference for making accurate and repeatable measurements of the real 
bone. The calibration disc measurements automatically and continuously 
calibrate the values calculated for bone density and thereby compensate 
for any short or long term drift in the x-ray detection electronics or 
other system variations, as well as differences in patient thickness. 
FIG. 6 shows schematically a plot of the function F=1n(L)-k*1n(H) for both 
the "bone" and "no bone" pulse pairs for a single traverse across the 
spine. It is important for this calibration to traverse across at least 
some portions of the patient having only flesh adjacent to the spine. 
(Using the notation defined above to be more precise, the calibration bone 
version of F, F(B) is equal to 1n(LB)-k*1n(HB) and the no calibration bone 
version of F, F(NB), is equal to 1n(L)-k*1n(H).) In these formulae, the 
symbol, 1n, indicates the natural logarithm function and the letter, k, is 
equal to the ratio of the attenuation coefficient of tissue for the LEL 
pulse to the attenuation coefficient of tissue for the HEL pulse. The 
values for H, L, HB, and LB used to calculate F in these formulae are the 
x-ray beam attenuation values measured by the main detector after 
corrections derived from the reference detector measurements have been 
applied. The reference corrections applied in this manner are given by the 
values for P0 as described above. 
The Wahner et al. publication cited demonstrates that the value of the 
functions F(B) and F(NB) will be a constant if there is no bone or 
bone-like material (or other high atomic number material) in the beam path 
through the patient. This is strictly true if k is constant across the 
scan line, independent of patient thickness, and can be made to hold in 
practice by measuring any dependence of k on patient thickness and using 
the corrected value of k to calculate the function F. In FIG. 6, the 
resulting constant values obtained when the beam is on either side of the 
spine (i.e. passing through portions of the body having flesh, without 
bone) are labeled Calibration Baseline and Normal Baseline corresponding 
to the plots of the functions F(B) and F(NB) respectively. The increase in 
the value of the function, F, over and above each baseline level when the 
x-ray beam scans across the spine has been shown to be directly 
proportional to the amount of bone or bone-like material in the path of 
the beam. 
In FIG. 6, the separation value 16 between the Normal Baseline and the 
Calibration Baseline can be calculated by finding the numerical average of 
the difference between F(B) and F(NB) for measurements made through 
portions of the body having only flesh on either side of the bone. This 
separation value is an important parameter in the operation of the 
densitometer because it calibrates the bone measurements in the spine 
directly against the known density of the bone-like calibration material 
which is used in the calibration disc measurements. Thus, the separation 
value corrects for both x-ray source drift, and for variations between 
different patients e.g. patient thickness, and other system variations. In 
FIG. 6, for example, suppose that a particular set of H/L measurements 
results in an F value (17 in FIG. 6) equal to 2.46 at the spine, and that 
the average separation value 16 at adjacent portions of the body has a 
value of 1.23. Then the bone density in the spine at the point where F 
equals 2.46 will be exactly (in this example) 2.00 times the of the 
bone-like calibration material used in the calibration disc. 
Using the separation value 16 as a calibration constant, the spine bone 
density can be calculated for all measurement pairs H/L recorded during 
the entire two dimensional raster scanning of the patient without the need 
for precise mechanical positioning of the x-ray beam or the calibration 
bone. The resulting two dimensional array of bone density values can be 
displayed as an image using readily available computer peripherals. 
Using the calibration disc to perform calibration measurements is the 
presently preferred method of calibrating the x-ray densitometer. Because 
calibration measurements are made numerous times for every scan line in 
the bone density image, one obtains assurance of proper calibration 
regardless of the location of the bone in the scan pattern. Other means of 
performing calibration measurements are possible and are comprehended by 
certain broader aspects of this invention. For example, if the pencil beam 
were brought to a rest at the end of each scan line or at the start or 
completion of the entire san a piece of bone-like material could be 
inserted in the beam and calibration data could be taken. Alternatively, 
the bone-like calibration material could be inserted into the beam on 
every alternate scan line. In another embodiment, a small amount of 
calibration material is placed under a part of the patient having only 
flesh without bone. 
After the baselines have been determined for each scan line and the bone 
density image has been created, it is also useful to detect automatically 
the outer boundary of the spine for each scan line. From these data it is 
possible to calculate a number of useful parameters. For example, the 
integral of the bone density values (expressed in units of grams per 
square centimeter) between two pre determined scan lines (usually chosen 
to correspond for spine measurements to a fixed number of vertebral 
bodies) is called the total bone mineral content (expressed in grams) and 
is one often quoted parameter. The integral of bone density values over 
the entire region of interest divided by the area of the bone as 
determined by the outer boundaries of the spine is the area averaged bone 
mineral density (expressed in grams per square centimeter) and is another 
frequently cited parameter. 
The means for determining baseline values for individual scan lines, the 
means for calculating the boundary of the bone in the bone density image, 
and other details of the bone density calculation are well known and are 
not novel to the present invention. The means for intercepting the x-ray 
beam with a bone-like calibration material and calculating two baselines 
at portions of the body having only flesh, to calibrate the system, has 
many advantages. 
FIG. 7 is a flow diagram illustrating the steps performed by computer 
system 13 to calculate bone mineral content of the spine while FIG. 7a 
diagrams the steps of the calibration routine, block 3 of FIG. 7 and FIG. 
7b diagrams the routine for calculation of bone mineral content. 
For calibration, step 3, the program examines the data and decides which 
points in the scan form part of the spine and which do not. In other 
words, it distinguishes "spine" points from "flesh" points. It achieves 
this very simply by setting respective thresholds for F(N) and F(NB) 
values and defines all values above the threshold as "spine" points and 
all values below the threshold as "flesh" points. 
The program then subtracts the F(NB) points that overlay flesh from the 
F(B) points that overlay flesh to calculate the separation value S.sub.L 
shown in FIG. 6 as distance 16. The average of the separation values 
averaged over all such subtracted pairs of points is then used as a 
calibration constant for Step 15 of FIG. 7 to calibrate the bone mineral 
content, see FIG. 7b. If the separation value changes by two percent, for 
example, from one day to another because of x-ray tube drift then the 
uncalibrated bone mineral content will change by two percent but the 
calibrated bone mineral content will remain constant. 
Most of the steps shown in FIG. 7 are self-explanatory but some additional 
explanation is helpful. After the separation value, calculated in Step 3, 
is subtracted from the F values for "bone", there is no longer any 
distinction between the F values for "bone" and "no bone" and these two 
sets of F values may be merged into a single set in order to increase the 
spatial resolution achieved by doubling the number of points used to 
create the bone image (Steps 4 and 5). 
In Step 7, the edges of the spinal vertebral bodies are found because FIG. 
7 illustrates the flow sequence for measuring bone mineral density of the 
spine. For measurements of other bones, such as the hip, a different edge 
detection routine would be used. Between Steps 8 and 9 and between Steps 
11 and 12, the operator is given an opportunity to modify the parameters 
calculated by the computer and to correct the parameters based on viewing 
the bone mineral image display. In Step 13 the operator decides which 
vertebral bodies will be included in the region of interest used to 
calculate bone mineral content and bone mineral density.