Non-invasive blood pressure sensor with motion artifact reduction and constant gain adjustment during pressure pulses

A non-invasive blood pressure sensor includes a first fluid filled sensing chamber having a diaphragm. A first transducer is fluidly coupled to the first sensing chamber to sense fluid pressure within the first chamber. A flexible body conformable wall surrounds the sensing chamber. The wall applies force to the artery while preventing pressure in a direction generally parallel to the artery from being applied to the sensing chamber. The flexible body conformable wall includes a second fluid filled chamber. A second transducer fluidly coupled to the second chamber senses fluid pressure within the second chamber. As varying pressure is applied to the artery pressure waveforms are sensed by the first transducer. Using output signals of the first and second transducers, the sensed pressure waveform data is analyzed to derive waveform parameters from which blood pressure values are derived. The effects of motion artifacts are reduced by the use of signals from both the first and second transducers. Part of the analysis of the waveform data includes the use of an adjusted gain that is substantially constant during pressure pulses, but that may vary from pulse to pulse. Signal values obtained from the second transducer are multiplied by the adjusted gain.

BACKGROUND OF THE INVENTION
 The present invention relates to systems and devices for measuring arterial
 blood pressure. In particular, the invention relates to an improved method
 and device for measuring arterial blood pressure in a non-invasive manner
 while reducing the effects of motion artifacts using a constant gain
 adjustment during pressure pulses.
 There has been a continuing need for devices which will measure blood
 pressure non-invasively, with accuracy comparable to invasive methods.
 Medwave, Inc. the assignee of the present invention, has developed
 non-invasive blood pressure measurement devices which are described in the
 following United States patents: U.S. Pat. No. 5,649,542 entitled
 CONTINUOUS NONINVASIVE BLOOD PRESSURE MONITORING SYSTEM; U.S. Pat. No.
 5,450,852 entitled CONTINUOUS NON-INVASIVE PRESSURE MONITORING SYSTEM;
 U.S. Pat. No. 5,640,964 entitled WRIST MOUNTED BLOOD PRESSURE SENSOR; U.S.
 Pat. No. 5,720,292 entitled BEAT ONSET DETECTOR; U.S. Pat. No. 5,738,103
 entitled SEGMENTED ESTIMATION METHOD; U.S. Pat. No. 5,722,414 entitled
 CONTINUOUS NON-INVASIVE BLOOD PRESSURE MONITORING SYSTEM; U.S. Pat. No.
 5,642,733 entitled BLOOD PRESSURE SENSOR LOCATOR; and U.S. Pat. No.
 5,797,850 entitled METHOD AND APATUS FOR CALCULATING BLOOD PRESSURE OF
 AN ARTERY. Further description of these devices is found in United States
 patent application Ser. No. 08/912,139 filed Aug. 15, 1997, entitled
 HAND-HELD NON-INVASIVE BLOOD PRESSURE MEASUREMENT DEVICE.
 As described in these patents and the pending patent application, the
 Medwave non-invasive blood pressure measurement device and method
 determines blood pressure by sensing pressure waveform data derived from
 an artery. A pressure sensing device includes a sensing chamber with a
 diaphragm which is positioned over the artery. A transducer coupled to the
 sensing chamber senses pressure within the chamber. A flexible body
 conformable wall is located adjacent to (and preferably surrounding) the
 sensing chamber. The wall is isolated from the sensing chamber and applies
 force to the artery while preventing pressure in a direction generally
 parallel to the artery from being applied to the sensing chamber.
 As varying pressure is applied to the artery by the sensing chamber,
 pressure waveforms are sensed by the transducer to produce sensed pressure
 waveform data. The varying pressure may be applied automatically in a
 predetermined pattern, or may be applied manually in a somewhat random
 fashion.
 The sensed pressure waveform data is analyzed to determine waveform
 parameters which relate to the shape of the sensed pressure waveforms. One
 or more blood pressure values are derived based upon the waveform
 parameters. The Medwave blood pressure measurement devices include both
 automated devices for continuously monitoring blood pressure (such as in a
 hospital setting) and hand-held devices which can be used by a physician,
 or by a patient when desired. These devices represent an important
 improvement in the field of non-invasive blood pressure measurement. Still
 further improvements, of course, are highly desirable.
 BRIEF SUMMARY OF THE INVENTION
 The present invention is an improvement to a non-invasive blood pressure
 sensing device and method of the type having a fluid filled sensing
 chamber and a flexible body conformable wall proximate to and isolated
 from the sensing chamber which applies force to the artery. The present
 invention is an improvement which minimizes effects of motion artifacts on
 the blood pressure measurement while maintaining accuracy of blood
 pressure readings using a constant gain adjustment during pressure pulses.
 In the present invention, the flexible body conformable wall includes a
 chamber which is separate from the sensing chamber. A first transducer
 senses pressure within the sensing chamber, while a second transducer
 senses pressure within the chamber which is a part of the flexible body
 conformable wall.
 The signals from the first and second transducers are processed and used to
 derive pressure waveform data from which blood pressure values are
 derived. Part of the processing includes the use of an adjusted gain that
 is substantially constant during heart beats, but that may vary from beat
 to beat. Signal values obtained from the second transducer are multiplied
 by the adjusted gain and then subtracted from the signal values obtained
 from the first transducer. The use of signals from both the first and the
 second transducers eliminates fluctuations in the signal from the first
 transducer which are the result of motion artifacts.

DETAILED DESCRIPTION
 FIG. 1 illustrates a hand held blood pressure measurement device being used
 to measure and display blood pressure within an underlying artery within
 wrist 12 of a patient. With device 10, a small amount of force is manually
 applied to the radial artery at the projection of the styloid process
 bone. As the force is manually applied, blood pressure waveforms are
 recorded and the corresponding hold down pressure which is being manually
 applied is also recorded. Using the pressure shape of the blood pressure,
 waveform parameters are generated. These parameters, along with universal
 coefficients, are used to calculate pressure values which then can be
 displayed.
 Blood pressure measurement device 10 includes main housing 14, display
 panel 16, on/off (power) and display select switch 20, sensor interface
 assembly 22, and connection plug 24.
 Housing 14 contains all of the electrical components of measurement device
 10. The diameter and length of housing 14 allow it to be easily held by
 the user (either medical personnel or the patient) during the measurement
 process. The hold down force is applied by applying force in an axial
 direction to wrist 12 which is transmitted from housing 14 to sensor
 interface assembly 22.
 Display panel 16 is preferably a liquid crystal display (LCD). In a
 preferred embodiment, display panel 16 simultaneously displays the
 following values based upon blood pressure measurements: systolic
 pressure, diastolic pressure, pulse rate, and mean blood pressure. Display
 panel 16 also preferably provides visual prompting for manually applying a
 varying hold down pressure.
 Power switch 20 is actuated to turn on power to the circuitry within
 housing 14. Timing circuitry within housing 14 automatically turns power
 off after a predetermined period of inactivity. Actuation of switch 20,
 after the unit is turned on, causes the display to indicate previous
 readings of blood pressure and pulse rate. In one embodiment there are ten
 memory locations for readings that can be displayed.
 Sensor interface assembly 22 is pivotally mounted to housing 14. As
 pressure is manually applied by moving housing 14 toward the artery, that
 force is transferred from housing 14 to sensor interface assembly 22.
 In operation, sensor interface assembly 22 is positioned over an artery
 such as the radial artery (as illustrated in FIG. 1). Alternatively,
 device 10 can be used in a number of other locations, such as on the
 temporal artery or the dorsalis pedis artery. The user then begins to
 apply force to the artery by applying axial force from housing 14 to
 sensor interface assembly 22. The force applied to the artery is swept in
 an increasing fashion so that pressure waveform data from a series of
 pulses are obtained with different amounts of force being applied. To
 achieve the desired pattern of variable force, user feedback is preferably
 provided with device 10.
 In a preferred embodiment, feedback is in the form of audible tones and/or
 movable bars on display 16 as shown in FIGS. 1A-1D. Top bar 16A is a
 pacing bar controlled by the microprocessor. Bottom bar 16B moves in
 response to the hold down pressure the user applies to the wrist through
 sensor interface assembly 22. As pressure is applied, bar 16A moves at a
 fixed rate. The user causes bottom bar 16B to move at approximately the
 same rate as top bar 16A by applying a steadily increasing force.
 The sequence of the measurement cycle is shown in FIGS. 1A-1D. First, the
 user presses power switch 20, which turns on the device 10. To take a
 reading, sensor interface assembly 22 is lightly pressed against a pulse
 locator (as illustrated in FIG. 1) so that bottom bar 16B remains under
 top bar 16A.
 Top bar 16A will start to move across display screen 16. As top bar 16A
 starts to move, the user must apply increasing pressure through device 10
 to the wrist so that bottom bar 16B tracks with the movement of top bar
 16A.
 FIG. 1B shows display 16 as top bar 16A has started to move from left to
 right and bottom bar 16B has not yet begun to track the movement of top
 bar 16A. FIG. 1C shows bars 16A and 16B as the process continues. Both
 bars are continuing to move from left to right across the bottom of the
 display 16. The amount of force required to keep bottom bar 16B underneath
 top bar 16A will increase as top bar 16A moves across display 16 from left
 to right.
 After a beep, the user can remove sensor interface assembly 22 from the
 wrist. At that point, top bar 16A returns to its left-most position, and
 bar 16B does not appear on the screen. This is shown in FIG. 1D. The user
 can then note the blood pressure reading. In a preferred embodiment
 illustrated in FIG. 1D, display 16 provides a digital readout of systolic,
 diastolic, and mean blood pressure, as well as pulse rate. An indication
 of memory location (by number) is also displayed.
 As soon as the reading is complete, device 10 is ready to take another
 reading. There is no need to clear display 16. Device 10 stores a
 predetermined number of previous readings (such as the last 10 readings).
 To review prior readings, power switch 20 is pressed. This causes a
 different reading from memory to be displayed on display 16.
 If a tone method is used as feedback, the user applies a force and each
 tone is modulated and has a higher pitch sound as the amplitude of the
 cardiac waveform increases. By listening to the tone, the user knows at
 what rate to apply the pressure to the artery. At the point of maximum
 energy transfer between the artery and sensor interface assembly 22, the
 cardiac pressure waveform reaches a peak amplitude and, therefore, the
 highest frequency tone is produced. As the user continues to apply higher
 pressure to the artery, the amplitude of the cardiac pressure waveform
 decreases, and therefore the frequency of the tone decreases. By listening
 to the tone, the user can perform a variable pressure sweep to measure
 pressure using device 10.
 Feedback to the user can be supplied in other ways as well. For example, an
 audible tone can be produced using a combination of frequency modulation
 and amplitude modulation. In other words, as the amplitude of the pressure
 waveform increases, both pitch (frequency) and amplitude (volume or
 loudness) of the tone will change.
 FIG. 2 is an electrical block diagram of device 10. Pressure transducers
 26A and 26B and nonvolatile memory 28 within sensor interface assembly 22
 are connected through connector 30 and connector 24 to circuitry within
 housing 14. Power supply circuit 31 includes switch 20, OR circuit 32,
 delay circuit 33, battery 34, reverse battery protection 35, integrated
 power switch 36, analog regulator 37, and voltage dividers 38A and 38D.
 The output of analog regulator 37 is electrical power which is used to
 energize analog circuitry, which includes amplifiers 40A and 40B, and
 analog-to-digital (A/D) converter 42. Integrated power switch 36 supplies
 power to all digital circuits, microprocessor 44, speaker 46, display
 panel 16 and associated display drive and memory circuitry 48.
 Microprocessor 44 includes digital signal processing circuitry 50, read
 only memory (ROM) and electrically erasable programmable read only memory
 (EEPROM) 52, random access memory (RAM) 54, timer circuitry 56, and
 input/output ports 58. A/D converter 42 may be integrated with
 microprocessor 44, while some of the memory may be external to
 microprocessor 44.
 Switch 20 is partially a monitoring pushbutton switch. Pressing switch 20
 causes OR circuit 32 to turn on integrated power switch 36. Integrated
 power switch 36 supplies power to microprocessor 44, which in turn latches
 on OR circuit 32. The turn off of the circuit is controlled by
 microprocessor 44 discontinuing a signal to OR circuit 32. This occurs
 through a fixed time of no activity.
 Transducers 26A and 26B sense pressure communicated within sensor interface
 assembly 22 and supply electrical signals to connector 30. In a preferred
 embodiment, transducers 26A and 26B are piezoresistive pressure
 transducers. Nonvolatile memory 28 stores offsets of transducers 26A and
 26B and other information such as sensor serial number. Nonvolatile memory
 28 is, in a preferred embodiment, an EEPROM.
 The outputs of transducers 26A and 26B are analog electrical signals
 representative of sensed pressure. These signals are amplified by
 amplifiers 40A and 40B and applied to inputs of A/D converter 42. The
 analog signals are converted to digital data and supplied to the digital
 signal processing circuitry 50 of microprocessor 44.
 Based upon the pressure data received, microprocessor 44 performs
 calculations to determine blood pressure values. Those calculations will
 be described in more detail later. As each pulse produces a cardiac
 waveform, microprocessor 44 determines a peak amplitude of the waveform.
 Microprocessor 44 controls display driver 48 to create bars 16A and 16B of
 FIGS. 1A-1D or drives speaker 46 to produce audible tones which vary as a
 function of the hold down pressure. The moving bars or audible tones guide
 the user in applying a variable force to the artery.
 When a measurement cycle has been completed, microprocessor 44 reorders the
 cardiac waveforms in increasing order of their corresponding hold down
 pressure and performs calculations to determine systolic pressure,
 diastolic pressure, mean blood pressure, and pulse rate. These values are
 displayed as shown in FIG. 1D. If switch 20 is pressed while
 microprocessor 44 is on, a signal is supplied through delay circuit 33 to
 microprocessor 44, causing it to toggle to a new pressure reading. The
 memory location of that pressure reading is also displayed, as shown in
 FIG. 1D.
 FIGS. 3A and 3B illustrate sensor interface assembly 22 in detail. Sensor
 interface assembly 22 includes top plate 150, upper cup 152, upper capture
 154, diaphragm capture 156, inner mounting ring 158, outer mounting ring
 160, side wall diaphragm 162, damping ring 164, inner diaphragm 166, and
 outer diaphragm 168.
 As shown in FIG. 3B, transducer 26A measures fluid pressure in fluid-filled
 sensor chamber A. Channels B, C, D, and E provide fluid pressure
 communications between transducer 26A and sensor chamber A. Transducer 26B
 measures fluid pressure in fluid-filled ring chamber F. Channels G and H
 provide fluid pressure communications between transducer 26B and ring
 chamber B. Connector 30 communicates with transducers 26A and 26B and
 non-volatile memory 28.
 FIG. 3B also shows how the sensor interface assembly 22 is pivotally
 connected to housing 14 by a ball 146 and socket 144 arrangement. The ball
 146 is pivotally mounted in socket 144. Because sensor interface assembly
 22 is pivotally coupled to stem 148 about a low pivot point. This permits
 sensor interface assembly 22 to be stably positioned above the underlying
 artery. In addition, the low pivot point enables the user to apply a more
 direct, uniform force on outer diaphragm 168. Thus, the hold down pressure
 manually applied by the user (through housing 14 and stem 148) is more
 uniformly applied to the anatomy above the underlying artery.
 Side wall diaphragm 162 and rings 158 and 160 define annular deformable
 ring chamber F coupled to ring 164. Side wall diaphragm 162 is preferably
 formed from a generally circular sheet of flexible material, such as
 polyurethane, and is filled with fluid. Diaphragm 162 has a hole sized to
 fit around the upper portion of upper capture 154. The outer edge portion
 of diaphragm 162 is trapped and held between outer ring 160 and top plate
 150. The inner edge portion of diaphragm 162 is trapped and supported
 between ring 158 and upper capture 154. Diaphragm 162 is made from a
 flexible material and is bulged outward when ring chamber F is filled with
 fluid. Ring chamber F is compressible and expandable in the vertical
 direction so as to be able to conform to the anatomy of the patient
 surrounding the underlying artery. As a result, the distance between top
 plate 150 and the patient's anatomy can vary around the periphery of side
 wall diaphragm 162 according to the contour of the patient's anatomy.
 Furthermore, because fluid is permitted to flow through and around chamber
 F, pressure is equalized around the patient's anatomy.
 Damping ring 164 generally consists of an annular compressible ring and is
 preferably formed from a foam rubber or other pulse dampening material
 such as open celled foam or closed cell foam. Ring 164 is centered about
 and positioned between side wall diaphragm 162 and diaphragms 166 and 168.
 Damping ring 164 is isolated from the fluid coupling medium within sensor
 chamber A. Because ring 164 is formed from a compressible material, ring
 164 absorbs and dampens forces in a direction parallel to the underlying
 artery which are exerted by the blood pressure pulses on sensor interface
 assembly 22 as the blood pressure pulse crosses sensor interface assembly
 22. Because bottom ring 164 is isolated from the fluid coupling medium in
 sensor chamber A, the forces absorbed or received by ring 164 cannot be
 transmitted to the fluid coupling medium. Instead, these forces are
 transmitted across ring 164 and side wall diaphragm 162 to top plate 150.
 Because this path is distinct and separate from the fluid coupling medium,
 sensor chamber A and the fluid coupling medium are isolated from these
 forces. In addition, ring 164 also presses tissue surrounding the artery
 to neutralize or offset forces exerted by the tissue.
 Upper diaphragm 166 is an annular sheet of flexible material having an
 inner diameter sized to fit around diaphragm capture 156. An inner portion
 of upper diaphragm 166 is trapped or captured (and preferably adhesively
 affixed) between the lip of diaphragm capture 156 and the bottom rim of
 upper capture 154.
 The intermediate portion of upper diaphragm 166 is adjacent to expansion
 cavity I and is isolated from ring 164 and ring chamber F. Upper diaphragm
 166 is permitted to initially move upward into expansion cavity I as ring
 chamber F, ring 164, and outer diaphragm 168 conform to the anatomy of the
 patient surrounding the underlying artery. As ring 164 is pressed against
 the anatomy of the patient surrounding the artery to neutralize or offset
 forces exerted by the tissue, outer diaphragm 168 is also pressed against
 the anatomy and the artery. However, because upper diaphragm 166 is
 permitted to roll into expansion cavity I, sensor chamber A does not
 experience a large volume decrease and a large corresponding pressure
 increase. Thus, sensor interface assembly 22 permits greater force to be
 applied to the anatomy of the patient through ring 164 to neutralize
 tissue surrounding the artery without causing a corresponding large change
 in pressure within sensor chamber A as the height of the side wall
 changes. As a result, sensor interface assembly 22 achieves more
 consistent and accurate blood pressure measurements.
 Outer diaphragm 168 is a generally circular sheet of flexible material
 capable of transmitting forces from an outer surface to fluid within
 sensor chamber A. Outer diaphragm 168 is coupled to inner diaphragm 166
 and is configured for being positioned over the anatomy of the patient
 above the underlying artery. Outer diaphragm sheet 168 includes non-active
 portion or skirt and an active central portion. The skirt constitutes the
 area of diaphragm 168 where inner diaphragm 166 is heat sealed or bonded
 to outer diaphragm 168.
 The active portion of outer diaphragm 168 is not bonded to inner diaphragm
 166, and is positioned below and within the inner diameter of ring 164.
 The active portion of outer diaphragm 168 is the active area of sensor
 interface assembly 22 which receives and transmits pulse pressure to
 transducer 26A.
 The coupling medium within sensor chamber A and passages B-E may consist of
 any fluid (gas or liquid) capable of transmitting pressure from diaphragm
 168 to transducer 26A. The fluid coupling medium interfaces between the
 active portion of outer diaphragm 168 and transducer 26A to transmit blood
 pressure pulses to transducer 26A. Because the fluid coupling medium is
 contained within sensor chamber A and passages B-E, which are isolated
 from the side wall of sensor interface assembly 22, the fluid coupling
 medium does not transmit blood pressure pulses parallel to the underlying
 artery, forces from the tissue surrounding the underlying artery and other
 forces absorbed by the side wall to transducer 26A. Forces parallel to the
 underlying artery are dampened by the compressible material of ring 164.
 As a result, sensor interface assembly 22 more accurately measures and
 detects arterial blood pressure.
 Sensor interface assembly 22 provides external measurements of blood
 pressure in an underlying artery. Because sensor interface assembly 22
 senses blood pressure non-invasively, blood pressure is measured at a
 lower cost and without medical risks. Because sensor interface assembly 22
 is relatively small compared to the larger cuffs used with oscillometric
 and auscultatory methods, sensor interface assembly 22 applies a hold down
 pressure to only a relatively small area above the underlying artery of
 the patient. Consequently, blood pressure measurements may be taken with
 less discomfort to the patient. Because sensor interface assembly 22 does
 not require inflation or deflation, faster, more frequent measurement3 may
 be taken.
 Furthermore, sensor interface assembly 22 better conforms to the anatomy of
 the patient so as to be more comfortable to the patient and so as to
 achieve more consistent and accurate blood pressure measurements. Because
 ring chamber F is deformable and filled with fluid, ring chamber F better
 conforms to the anatomy of the patient and equalizes pressure applied to
 the patient's anatomy. Because ring 164 is compressible and because outer
 diaphragm 168 is flexible and is permitted to bow or deform inwardly, ring
 164 and outer diaphragm 168 also better conform to the anatomy of the
 patient. At the same time, however, sensor interface assembly 22 does not
 experience a large sudden increase in pressure in sensor chamber A as ring
 164 and outer diaphragm 168 are pressed against the anatomy of the
 patient. Ring chamber F and ring 164 apply force to the anatomy of the
 patient to neutralize the forces exerted by tissue surrounding the
 underlying artery. Because ring chamber F and ring 164 are both
 compressible in height, the height of the side wall decreases as the side
 wall is pressed against the patient. Diaphragms 166 and 168 arc also
 conformable. However, because the intermediate portion of inner diaphragm
 166 is permitted to move upward into expansion cavity I, sensor chamber A
 does not experience a large volume decrease and a corresponding large
 pressure increase. Thus, the side wall is able to apply a greater force to
 the anatomy of the patient without causing a corresponding large,
 error-producing increase in pressure within sensor chamber A due to the
 change in height of the side wall and the change in shape of outer
 diaphragm 168.
 At the same time, sensor interface assembly 22 permits accurate and
 consistent calculation of blood pressure. Because of the large sensing
 area through which blood pressure pulses may be transmitted to transducer
 26A, sensor interface assembly 22 is not as dependent upon accurate
 positioning of the active portion of outer diaphragm 168 over the
 underlying artery. Thus, sensor interface assembly 22 is more tolerant to
 patient movement as measurements are being taken.
 Moreover, sensor interface assembly 22 achieves a zero pressure gradient
 across the active face of the sensor, achieves a zero pressure gradient
 between the transducer and the underlying artery, attenuates or dampens
 pressure pulses that are parallel to the sensing surface of the sensor,
 and neutralizes forces of the tissue surrounding the underlying artery.
 Sensor interface assembly 22 contacts and applies force to the anatomy of
 the patient across the skirt and the active portion of outer diaphragm
 168. However, the pressure within sensor chamber A is substantially equal
 to the pressure applied across the active portion of outer diaphragm 168.
 The remaining force applied by sensor interface assembly 22 across the
 skirt, which neutralizes or offsets forces exerted by the tissue
 surrounding the underlying artery, is transferred through the side wall
 (ring 164 and ring chamber F) to top plate 150. As a result, the geometry
 and construction of sensor interface assembly 22 provides the proper ratio
 of pressures between the skirt and the active portion of outer diaphragm
 168 to neutralize tissue surrounding the underlying artery and to
 accurately measure the blood pressure of the artery. In addition, because
 the fluid coupling medium within sensor chamber A is isolated from the
 side wall, pressure pulses parallel to the underlying artery, forces from
 tissue surrounding the underlying artery, and other forces absorbed by the
 side wall are not transmitted through the fluid coupling medium to
 transducer 26A. Consequently, sensor interface assembly 22 also achieves a
 zero pressure gradient between transducer 26A and the underlying artery.
 Blood pressure measuring device 10 determines blood pressure values from
 the sensed waveform pressure amplitudes sensed by sensor interface
 assembly 22 and from other parameters derived from the pressure amplitudes
 using a stored set of coefficients. A pressure amplitude is determined at
 each sample point.
 Device 10 calculates a systolic blood pressure value (S), a mean blood
 pressure value (M) and a diastolic blood pressure value (D) based upon the
 following formulas:
EQU M=F.sub.m (P.sub.1.sup.m, . . . , P.sub.n.sup.m, C.sub.1.sup.m, . . .
 C.sub.n.sup.m)
EQU S=F.sub.s (P.sub.1.sup.s, . . . , P.sub.n.sup.s, C.sub.1.sup.s, . . .
 C.sub.n.sup.s)
EQU D=F.sub.d (P.sub.1.sup.d, . . . P.sub.n.sup.d, C.sub.1.sup.d, . . .
 C.sub.n.sup.d)
 wherein F.sub.m, F.sub.s, F.sub.d are linear or non-linear functions,
 P.sub.1.sup.m, P.sub.1.sup.s, P.sub.1.sup.d, . . . P.sub.n.sup.m,
 P.sub.n.sup.s, P.sub.n.sup.d are parameters derived from waveform pressure
 amplitudes, and C.sub.1.sup.m, C.sub.1.sup.s, C.sub.1.sup.d, . . . ,
 C.sub.n.sup.m, C.sub.n.sup.s, C.sub.n.sup.d are coefficients obtained
 during training processes based upon clinical data.
 In particular, device 10 calculates a systolic blood pressure value (S), a
 mean blood pressure value (M), a diastolic blood pressure value (D) based
 upon the following formulas:
EQU M=C.sub.1.sup.m P.sub.1.sup.m +C.sub.2.sup.m P.sub.2.sup.m + . . .
 +C.sub.n.sup.m P.sub.n.sup.m
EQU S=C.sub.1.sup.s P.sub.1.sup.s +C.sub.2.sup.s P.sub.2.sup.s + . . .
 +C.sup.n.sup.s P.sub.n.sup.s
EQU D=C.sub.1.sup.d P.sub.1.sup.d +C.sub.2.sup.d P.sub.2.sup.d + . . .
 +C.sub.n.sup.d P.sub.n.sup.d
 wherein P.sub.1.sup.m, P.sub.1.sup.s, P.sub.1.sup.d . . . P.sub.n.sup.m,
 P.sub.n.sup.s, P.sub.n.sup.d are parameters derived from waveform pressure
 amplitudes. Such parameters may be calculated from shape characteristics
 of the waveform or parameters calculated from functions such as curves
 based upon relationships between particular points of several waveforms.
 The parameters may be further based upon hold down pressure values and
 time periods between particular points on the waveforms. The values
 C.sub.1.sup.m, C.sub.1.sup.s, C.sub.1.sup.d . . . C.sub.n.sup.m,
 C.sub.n.sup.s, C.sub.n.sup.d are coefficients obtained during training
 processes based upon clinical data.
 In addition, the pulse rate (PR) may also be determined using the formula:
 ##EQU1##
 To determine the pulse rate, four individual waveforms, or beats, are
 sensed and are time averaged to determine the pulse rate. Preferably, the
 waveforms used to determine pulse rates include the waveform having the
 largest maximum pressure amplitude, the two waveforms prior to the
 waveform having the largest maximum pressure amplitude and the waveform
 succeeding the waveform having the largest maximum pressure amplitude.
 Once the four waveforms are identified, the pulse rate of each waveform is
 determined. The sum of the pulse rate of the four waveforms is then
 divided by four to calculate pulse rate PR. The pulse rate (PR) for each
 waveform is based upon the following formula:
 ##EQU2##
 FIG. 4 illustrates a sample series of waveforms exhibited by the underlying
 artery as a varying pressure is applied over time. The vertical scale
 indicates pressure in mmHg while the horizontal scale indicates individual
 sample points at which the blood pressure values exerted by the pulse are
 measured over time. In the preferred embodiment, transducers 26A and 26B
 produce continuous electrical signals representing waveform pressures
 which are sampled 128 times per second.
 In the preferred embodiment, the hold down pressure applied to sensor
 interface assembly 22 is swept over a preselected range of increasing hold
 down pressures. Preferably, the sweep range of hold down pressures
 typically is begun at approximately 10 mmHg. The hold down pressure is
 then steadily increased (under the prompting or guidance from the audible
 or visual feedback) until two individual waveforms are sensed following
 the sensed waveform having the largest pressure amplitude. Preferably,
 each sweep range extends between an initial hold down pressure of about 10
 mmHg and a final hold down pressure of approximately 150% of the mean hold
 down pressure of the waveform having the largest maximum pressure
 amplitude during the previous sweep.
 FIG. 4 shows the signals 400 and 410 from transducers 26A and 26B,
 respectively, as sensor interface assembly 22 is pressed against the
 artery. Signal 400 is representative of pressure in sensor chamber A.
 Signal 410 represents the pressure in ring chamber F as sensed by
 transducer 26B. Signal 420 is representative of pressure in ring chamber F
 after applying a proper gain and offset. Signal 410 is calibrated to match
 signal 400. This gain and offset adjustment can take place in an initial
 phase of a pressure measurement. This gain and offset adjustment can also
 take place on a continual basis or at any other phase of a pressure sweep.
 A least square fit can be used to find the best fit of curves 400 and 410
 so as to get the best gain and offset adjustment.
 At multiple places during the sweep, signals 400 and 420 are affected by
 patient movement or (in the case of a hand-held blood pressure unit)
 operator movement as the sweep is performed. These inflections or motion
 artifacts MA show noise that needs to be taken out of the system in order
 to measure blood pressure. Signal 430 has most of the noise taken out of
 signal 400, and is referred to as a "clean" signal. Signal 430 contains
 pressure pulse waveforms 500 from the movement of the arterial walls as
 sensor interface assembly 22 is pressed against the arterial wall.
 FIG. 5 shows a summary of a preferred process for calculating "clean"
 signal values, which are the values used to construct curve 430 shown in
 FIG. 4. The first step shown in FIG. 5 is to obtain digital data samples
 from sensor chamber A and ring chamber F. (Block 450). The samples from
 sensor chamber A are represented by "main(t)", and the samples obtained
 from ring chamber F are represented by "ring(t)", where "t" represents the
 time at which each sample was taken. The digital data samples from sensor
 chamber A (i.e., main(t)) are represented graphically in FIG. 4 as curve
 400. The digital data samples from ring chamber F (i.e., ring(t)) are
 represented graphically in FIG. 4 as curve 410.
 As digital data samples are obtained, they are supplied to digital signal
 processing circuitry 50 of microprocessor 44. (Block 452). Preferably,
 digital signal processing circuitry 50 determines the variable relative
 gain (K(t)) between the main(t) values and the ring(t) values at each
 sample point. (Block 454). The variable relative gain between sample
 points may be calculated while digital data samples are being obtained. In
 contrast, using the least square fit method typically requires that all of
 the samples be obtained for a particular pressure sweep prior to
 calculating the variable relative gain.
 The ring(t) values are multiplied by the gain coefficients K(t) to obtain
 the values for signal 420. In a preferred embodiment, only a subset of the
 gain coefficients are used in adjusting curve 410 to obtain the values for
 curve 420. Specifically, only the gain coefficients with time values t
 that correspond to the beginning of each pressure pulse waveform or beat
 are used. These values are then held constant during the beat to generate
 an adjusted gain 472. FIG. 6 shows a graph of the variable gain
 coefficients 470 and the adjusted gain 472 versus time. In alternative
 preferred embodiments, the abrupt step changes in gain between beats in
 curve 472 are eliminated, and the transitions in gain between beats are
 made smooth. By smoothing the transitions in curve 472, the gain still
 remains substantially constant during each beat.
 The digital data samples ring(t) are preferably multiplied by the adjusted
 gain 472 to obtain the values for curve 420. The adjusted gain 472, rather
 than gain 470, is used in adjusting curve 410, because, in minimizing the
 error between curve 400 and curve 410, the variable gain method views the
 pressure pulse waveforms on signal 400 as errors that need to be
 minimized. However, the pressure pulse waveforms are actually signal
 values, and not errors.
 The next step in the process for generating "clean" signal values (curve
 430) is to identify the time at the beginning of each beat, or the beat
 onset (t.sub.0), for each beat in curve 400. (Block 456).
 Commonly-assigned U.S. Patent No. 5,720,292, entitled "BEAT ONSET
 DETECTOR", discloses a preferred method for detecting the onset of heart
 beats. The variable relative gain (K(t)) corresponding to the time of each
 beat onset is identified. (Block 458). Lastly, the "clean" signal values
 (curve 430) are obtained for each sample point using the equation:
EQU clean(t)=main(t)-K(t.sub.0).times.ring(t)
 where:
 clean(t) represents the digital data values for curve 430;
 main(t) represents the values for curve 400;
 ring(t) represents the values for curve 410; and
 K(t.sub.0) represents the variable relative gain at the beat onset for each
 beat. (Block 460).
 It has been determined that keeping the gain constant during each beat
 (i.e., using the adjusted gain 472 rather than gain 470) in calculating
 the values for curve 430 helps to minimize the distortion of the shape of
 each "clean" beat 500. Minimizing the distortion of the shape of each beat
 is important because, as discussed below, various parameters may be
 obtained based on the shape of these beats. For example, one parameter
 that is affected by the gain adjustment is the tail segment length
 parameter, which is illustrated in FIG. 7.
 FIG. 7 shows waveform 474, which is a normalized graph of a pressure
 waveform taken from curve 400, and also shows waveform 476, which is a
 normalized graph of one of clean pressure waveforms 500 taken from curve
 430, where curve 430 was calculated using the variable gain coefficients
 470 shown in FIG. 6. The tail segment length parameter represents the
 length of the flat portion at the end of each pressure waveform. For
 pressure waveform 474, the tail segment length parameter is represented by
 L.sub.1, and for pressure waveform 476, the tail segment length parameter
 is represented by L.sub.2. The tail segment length parameter is a
 distortion parameter taken from the pressure waveform occurring before the
 pressure waveform with the maximum amplitude.
 As shown in FIG. 7, when the variable gain coefficients 470 are used in
 calculating curve 430, the tail of the pressure waveform 476 becomes
 longer than the tail of pressure waveform 474. When the tail segment
 length parameter is used in calculating diastolic blood pressure, it has
 been determined that the increased value of the tail segment length
 parameter causes the calculated diastolic blood pressure to be slightly
 lower than readings obtained from an arterial line. The tail segment
 length parameter is over estimated due to the fact that the variable gain
 is adapting every sample and is thereby causing the end of each pressure
 waveform to be a little flatter and longer.
 By using adjusted gain 472 rather than gain 470 in calculating curve 430, a
 more accurate tail segment length parameter is obtained. FIG. 8 shows a
 graph of waveform 478, which is a normalized graph of a pressure waveform
 obtained directly from a patient's artery. As can be seen in FIG. 8, the
 tail segment length parameter of waveform 478 is virtually identical to
 the tail segment length parameter L, of waveform 474. By using the tail
 segment length parameter from waveform 478 in calculating diastolic blood
 pressure, the calculated blood pressure value more closely approaches the
 value obtained from the arterial line method. More accurate values are
 also obtained for systolic and mean blood pressure when parameters are
 derived from waveform 478 rather than waveform 476.
 Signal 430 represents blood pressure pulses that can be used to obtain
 shape and amplitude information to calculate blood pressure. Signal 400
 can be used to obtain additional information such as hold down pressure
 that is also used to calculate pressure.
 As can be observed in FIG. 4, when noise causes signal 400 to sweep in a
 non-uniform movement, it may be required to reorder the beats in order of
 increasing hold down pressure in order to calculate blood pressure.
 Based upon sensed and sampled pressure waveform signals or data produced by
 transducers 26A and 26B during each sweep of hold down pressures,
 microprocessor 44 derives preselected parameters for calculating blood
 pressure values from the derived parameters and a stored set of
 coefficients. As indicated in FIG. 4, parameters may be derived directly
 from the absolute waveform pressures which vary as hold down pressure is
 varied over time. Such parameters may be derived from the shape of the
 waveforms including a particular waveform's slope, absolute pressure at a
 selected sample point, a rise time to a selected sample point on a
 waveform, and the hold down pressures corresponding to a particular sample
 point on a waveform. As can be appreciated, any of a variety of parameters
 may be derived from the absolute waveform pressures shown in FIG. 4.
 Parameters may further be based upon particular points or functions of the
 sample points.
 A preferred process of calculating pressure using shape, amplitude, and
 hold down is described in commonly-assigned U.S. patent application Ser.
 No. 08/912,139, filed Aug. 15, 1997, entitled HAND-HELD NON-INVASIVE BLOOD
 PRESSURE MEASUREMENT DEVICE, and U.S. Pat. No. 5,797,850, entitled METHOD
 AND APATUS FOR CALCULATING BLOOD PRESSURE OF AN ARTERY, which are
 incorporated by reference.
 In preferred embodiments of the present invention, the waveform analysis
 described in U.S. Pat. No. 5,738,103 entitled "Segmented Estimation
 Method" and U.S. Pat. No. 5,720,292 entitled "Beat Onset Detector" are
 also used.
 Although the present invention has been described with reference to
 preferred embodiments, workers skilled in the art will recognize that
 changes may be made in form and detail without departing from the spirit
 and scope of the invention. For example, although the determination of
 pressure values based upon waveform parameters has been described using
 linear equations and stored coefficients, other methods using non-linear
 equations, look-up tables, fuzzy logic and neural networks also can be
 used in accordance with the present invention.
 In other embodiments, algorithms can be used that compensate for a
 nonlinear hold down pressure sweep. This is accomplished by recording hold
 down pressure and pulse shape, so that the operation does not perform a
 linear sweep. A linear sweep can be constructed as long as there are
 several pulse shapes recorded over the range of the sweep, regardless of
 the order they are recorded.