text: As is well known in the art, in medical diagnosis and treatment, it is often desirable to quantitatively measure over time the respiratory air volume or pulmonary ventilation. This has conventionally been done by having the patient or subject breathe into a mouthpiece connected to a flow rate measuring device. Flow rate is then integrated to provide air volume change.
  There are, however, several drawbacks and disadvantages associated with employing a mouthpiece. A mouthpiece is difficult to use for long term subject monitoring, especially for ill, sleeping or anesthetized subjects. Further, it is uncomfortable for the subject, tends to restrict breathing, and is generally inconvenient for the physician or technician to use.
  As is also well known in the art, there are qualitative respiration monitors available that do not require a mouthpiece. Illustrative are the systems disclosed in U.S. Pat. Nos. 3,831,586 and 4,033,332. Although the noted systems eliminate most of the disadvantages associated with a mouthpiece, the systems do not, in general, provide an accurate measurement of air volume. Further, the systems are typically only used to signal an attendant when a subject's breathing activity changes sharply or stops.
  Another means for quantitatively measuring respiratory or lung volume is to measure the change in size of the rib cage and abdomen, as it is well known that lung volume is a function of these two parameters. A number of systems and devices have been employed to measure the change in size (i.e. A circumference) of the rib cage and abdomen, including mercury in rubber strain gauges, pneumobelts, magnetometers, and respiratory inductive plethysmograph (RIP) belts, see, e.g., D. L. Wade, “Movements of the Thoracic Cage and Diaphragm in Respiration”, J. Physiol., pp. 124-193 (1954), Mead, et al, “Pulmonary Ventilation Measured from Body Surface Movements”, Science, pp. 196, 1383-1384 (1967).
  In practice, respiratory magnetometers and RIP belts are primarily used to measure the change in size of the rib cage and abdomen. As is well known in the art, respiratory magnetometers consist of tuned pairs of electromagnetic coils or magnetometers; one coil being adapted to transmit a specific high frequency AC electromagnetic field (i.e. transducer) and the other coil (i.e. receiver) being adapted to receive the field. To measure the anteroposterior diameter of the rib cage, a first coil, e.g., transducer, is typically placed over the sternum at the level of the 4th intercostal space and the second coil (of the pair) is placed over the spine at the same level. To measure the anteroposterior diameter of the abdomen, a third coil is typically placed on the abdomen at the level of the umbilicus and a fourth coil (of the pair) is placed over the spine at the same level.
  Over the operational range of distances, the output voltage is linearly related to the distance between a pair of coils; provided, the axes of the coils or magnetometers remain parallel to each other. As rotation of the axes can change the voltage, the transducer and receiver coils must be secured to the skin in a parallel fashion and rotation due to the motion of underlying soft tissue that must be minimized.
  A potential limitation of the use of such coils or magnetometers is presented in environments that contain large metal structures or electric motors. Such devices produce extraneous electromagnetic fields and consequently affect the magnetometer voltage output.
  RIP belts consist of two loops of wire that are coiled and sewed into an elastic belt. To measure changes in cross-sectional areas of the rib cage and abdomen, one belt is secured around the mid-thorax and a second belt is placed around the mid-abdomen.
  The voltage change from the belts is generally linearly related to changes in the enclosed cross-sectional area. When the RIP belts are operated in the DC-coupled mode, they can detect shifts in chest wall dimensions, e.g. a change of FRC. However, the AC-coupled mode is typically preferred for tidal volume measurements.
  For quantitative measurements, RIP uses a “two-degrees-of-freedom” model to assess changes in perimeters (i.e. cross-sectional area) of the rib cage and abdomen. Since the first rib and adjacent structures of the neck are relatively immobile, the moveable components of the thoracic cavity are taken to be the anterior and lateral walls of the rib cage and the abdomen. Changes in volume of the thoracic cavity will then be reflected by displacements of the rib cage and abdomen.
  Displacement (i.e. motion) of the rib cage can be directly assessed. Diaphragm displacement cannot be measured directly, but since the abdominal contents are essentially incompressible, caudal motion of the diaphragm relative to the pelvis and the volume it displaces is reflected by outward movement of the anterolateral abdominal wall.
  The “two-degrees-of-freedom” model embraced by most in the field holds that the volume displacement of the respiratory system, i.e. tidal volume (VT), is equal to the sum of the volume displacements of the rib cage and abdomen, i.e.VT=αRC+βAb  Eq. 1where:RC and Ab represent linear displacements of the rib cage and abdomen, respectively; and α and β represent volume-motion coefficients.
  As is well known in the art, RC and Ab linear displacements are converted to RC and Ab volume displacements when multiplied by the α and β volume-motion coefficients.
  It is well established that the use of the noted “two-degrees-of-freedom” model can provide an estimate of VT that is within 10% accuracy of ventilation measured at the mouth; provided, the subject is confined to one body position.
  Two different approaches primarily used for determining the necessary volume-motion coefficients of the rib cage and abdomen are the isovolume technique and the multiple linear regression technique. In the isovolume technique, the subject first performs an isovolume maneuver, shifting volume back and forth between the rib cage and abdominal compartments while holding the glottis closed, whereby there is no net volume change of the system. Since VT equals zero, Equation 1 can be modified as follows:RC=(−β/α)Ab  Eq. 2
  On a graph of rib cage and abdomen signals, the slope of the isovolume line is equal to the ratio −β/α.
  In practice, the gains of the rib cage and abdomen signals are often adjusted, whereby the slope of the isovolume line equals one. The rib cage and abdomen displacements are thus equal for any volume change. The two signals can then be directly summed to provide volume.
  The isovolume method is based on the assumptions that displacements of the surfaces of the rib cage and abdomen are representatively sampled at the measured location, and are similar during isovolume efforts and spontaneous breathing. Since volume-motion coefficients change with posture, the isovolume calibration must be repeated in each body position.
  Computer-assisted regression techniques, such as multiple linear regression, are used to determine volume-motion coefficients by solving a matrix of multiple simultaneous equations of changes in chest wall dimensions and lung volume. An advantage of these techniques is that no special calibration maneuver is required to generate volume-motion coefficients.
  A limitation of any approach that uses chest wall motion to assess ventilation is, however, that the overall volume change of the chest wall being measured includes not only changes in lung volume, but also blood volume shifts into and out of the thoracoabdominal cavity. This can occur when the respiratory system is subjected to large pressure changes, or with changes in posture (e.g. between supine and upright position).
  Another limitation is related to distortion that can occur within the rib cage or abdomen (e.g. between the upper and lower rib cage or between the lower transverse and AP rib cage).
  As is well known in the art, the accuracy of “two-degrees-of-freedom” model and, hence, methods employing same to determine volume-motion coefficients of the rib cage and abdomen, is further limited by virtue of changes in spinal flexion that can accompany changes in posture. Indeed, it has been found that VT can be over or under-estimated by as much as 50% of the vital capacity with spinal flexion and extension, see McCool, et al., “Estimates of Ventilation From Body Surface Measurements in Unrestrained Subjects”, J. Appl. Physiol., vol. 61, pp. 1114-1119 (1986); and Paek, et al., “Postural Effects on Measurements of Tidal Volume From Body Surface Displacements”, J. Appl. Physiol., vol. 68, pp. 2482-2487 (1990).
  There are two major causes that contribute to the noted “two-degrees-of-freedom” model error(s) and, hence, limitation. A first contributing cause of the error is due to the substantial displacement of the summed rib cage and abdomen signals that occurs with isovolume spinal flexion and extension or pelvic rotation, which is illustrated in FIG. 1.
  These shifts are a consequence of conservation of volume. As one of the thoracoabdominal boundaries is pushed in, another must be pushed out.
  The second contributing cause of the error is due to posturally-induced changes in volume-motion coefficients. With isovolume spinal flexion, the rib cage comes down with respect to the pelvis and the axial dimension of the anterior abdominal wall becomes smaller. Therefore, less abdominal cavity is bordered by the anterior abdominal wall.
  With a smaller anterior abdominal wall surface to displace, a given volume displacement of the abdominal compartment would be accompanied by a greater outward displacement of the anterior abdominal wall. The abdominal volume-motion coefficient would accordingly be reduced.
  It has, however, been found that the addition of a measure of the axial motion of the chest wall, i.e. changes in the distance between the xiphoid and the pubic symphysis (Xi), provides a third degree of freedom, which, when employed to determine VT can reduce the posture related error associated with the “two-degrees-of-freedom” model to within 15% of that measured by spirometry, see Paek, et al., “Postural Effects on Measurements of Tidal Volume From Body Surface Displacements”, J. Appl. Physiol., vol. 68, pp. 2482-2487 (1990); and Smith, et al. “Three Degree of Freedom Description of Movement of the Human Chest Wall”, J. Appl. Physiol., Vol. 60, pp. 928-934 (1986).
  Smith, et al. proposed the following “three-degrees-of-freedom” model to determine tidal volume, i.e.VT=αRC+βAb+γXi  Eq. 3where:RC and Ab represent linear displacements of the rib cage and abdomen, respectively;Xi represents the Δ distance between the xiphoid and the pubic symphysis; andα, β and γ represent volume-motion coefficients for RC, Ab and Xi.
  Referring to FIG. 2, there are shown graphical illustrations of % error of estimated VT determined with one, two and three degrees of freedom models (i.e. x-axis) during various postural movements or maneuvers. It can be seen that the use of a “three-degrees-of-freedom” model incorporating the third independent variable, i.e. the Δ distance between the xiphoid and the pubic symphysis (“Xi”), enhances the accuracy with which volume is estimated from body surface motion in those maneuvers that incorporate changes in spinal attitude.
  There are, however, similarly several drawbacks and disadvantages associated with the “three-degrees-of-freedom” model. A major drawback is that the “three-degrees-of-freedom” model reflected in Eq. 3 above is still limited in accuracy to about 15% of actual ventilation in individuals who are doing freely moving postural tasks, such as bending, sitting or standing, due to spinal flexion.
  It would thus be desirable to provide an improved method and associated system for determining tidal volume (or pulmonary ventilation) that substantially reduces or eliminates the drawbacks and disadvantages associated with conventional methods and systems that are employed to determine pulmonary ventilation.
  It is therefore an object of the present invention to provide noninvasive methods and associated systems for determining pulmonary ventilation that substantially reduce or eliminate the drawbacks and disadvantages associated with conventional methods and systems for determining pulmonary ventilation.
  It is another object of the invention to provide noninvasive methods and associated systems for determining pulmonary ventilation that substantially reduce the accuracy errors associated with conventional two-degrees and three-degrees of freedom tidal volume models.
  It is another object of the invention to provide noninvasive methods and associated systems for determining pulmonary ventilation that can be readily employed to measure pulmonary ventilation in different postures when awake and during sleep.
  It is another object of the invention to provide noninvasive methods and associated systems for determining pulmonary ventilation that can be readily employed to accurately detect respiratory abnormalities.
  It is yet another object of the invention to provide noninvasive methods and associated systems for determining pulmonary ventilation that can be readily employed to accurately detect respiratory events, such as apneas and hypopneas, during sleep.