CA2054941C - Cardiac pacemaker - Google Patents
Cardiac pacemaker Download PDFInfo
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- CA2054941C CA2054941C CA002054941A CA2054941A CA2054941C CA 2054941 C CA2054941 C CA 2054941C CA 002054941 A CA002054941 A CA 002054941A CA 2054941 A CA2054941 A CA 2054941A CA 2054941 C CA2054941 C CA 2054941C
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- 230000000747 cardiac effect Effects 0.000 title claims description 17
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Classifications
-
- A—HUMAN NECESSITIES
- A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
- A61N—ELECTROTHERAPY; MAGNETOTHERAPY; RADIATION THERAPY; ULTRASOUND THERAPY
- A61N1/00—Electrotherapy; Circuits therefor
- A61N1/18—Applying electric currents by contact electrodes
- A61N1/32—Applying electric currents by contact electrodes alternating or intermittent currents
- A61N1/36—Applying electric currents by contact electrodes alternating or intermittent currents for stimulation
- A61N1/362—Heart stimulators
- A61N1/365—Heart stimulators controlled by a physiological parameter, e.g. heart potential
- A61N1/36514—Heart stimulators controlled by a physiological parameter, e.g. heart potential controlled by a physiological quantity other than heart potential, e.g. blood pressure
- A61N1/36542—Heart stimulators controlled by a physiological parameter, e.g. heart potential controlled by a physiological quantity other than heart potential, e.g. blood pressure controlled by body motion, e.g. acceleration
Abstract
An accelerometer based rate adaptive pacemaker. An acceleration sensor generates an acceleration output signal corresponding to sensed acceleration of a patient's body. Apparatus for providing a sensor determined interval is coupled to the acceleration sensor and the sensor determined interval is proportional to the acceleration output signal. Apparatus for determining an actual pacing interval is coupled to an apparatus for determining a generated pacing interval. The generated pacing interval is a function of both the sensor determined interval and the actual pacing interval.
Description
ACCELEROMETER BASED RATE ADAPTIVE CARDIAC PACEMAKER
BACKGROUND OF THE INVENTION
I. Field of the Invention This invention relates to signal processing and rate modulation in automatic cardiac pacers and, more particularly, to a rate adaptive pacerwhich provides pacing signals to the heart based on a root mean square approximation of band limited, low frequency, low level acceleration signals generated from body motion.
II Description of the Prior Art Activity sensing rate-responsive pacemakers have been developed which use body vibration during physical activity as an indicator of the need to alter pacing rate. One such prior art device is described in US Patent No. 4,428,378 to Anderson, et al., entitled "Rate Adaptive Pacer". The Anderson, et al. pacemaker operates by responding to a variety of mechanical forces both internal and external to a pacemaker patient. Because the Anderson, et al. pacemaker responds to forces applied to the patient's body, it must necessarily rely on the outside environment to provide a mass to generate the measured forces. Such masses as provided by the outside environment are subject to wide variations. Yet another deficiency of such prior art sensors is that they can measure only changes in pressure, not constant pressure. The present invention provides an activity sensing rate-responsive pacemaker including an accelerometer which contains a captive, well defined, constant reference mass.
Prior art rate adaptive pacemakers which are dependent on mechanical force sensors have been found to be relatively insensitive to different levels of exertion, particularly with respect to providing appropriate rate changes in response to a patient ascending stairs and walking up inclines. Such p~~~cf ~~C~r -~-are also susceptible to the effects of extraneous vibration such as that encountered in various~forms of transport. One such study is reported by C.P. Lau, et al. in an article entitled "Selective Vibration Sensing: A New Concept for Activity - Sensing Rate-Responsive Pacing," P. 1299, PACE. Vol. II, Sept. 1988.
BUMMARY OF THE INVENTION
An accelerometer based rate adaptive pacemaker is provided.
Means for sensing acceleration is included wherein the acceleration sensing means provides an acceleration output signal corresponding to the sensed acceleration. Means for a sensor determined interval is coupled to the acceleration sensing means wherein the sensor determined interval is proportional to the acceleration output signal. Means for determining an actual pacing interval provides the actual pacing interval to a means for determining a pacing rate. The pacing rate means is, also, coupled to the sensor determined interval means. The pacing rate is a function of both the sensor determined interval and the actual pacing interval.
In one aspect of the invention, the programmed pacing interval is decreased if the actual pacing interval exceeds the sensor determined interval.
In another aspect of the invention, a root mean square (RMS) approximation is applied to a sensor output to yield a sensor rectified average value which is used for pacing rate modulation.
In yet another aspect of the invention, analog means are provided Eor amplifying and band limiting a low frequency, low level accelerometer signal with very low noise.
In still another aspect of the invention, a method is provided for converting signal energy into a pacing rate.
Other objects, features and advantages of the present invention will become apparent to those skilled in the art through the Description of the Preferred Embodiment, Claims, and drawings herein wherein like numerals refer to like elements.
BRIEF DESCRIPTION OF THE DRA'OPIN(38 Figure 1 is a block diagram of a cardiac pacemaker as provided by one embodiment of the invention.
Figure 2 is a block diagram of an analog signal processing network as employed in one embodiment of the invention.
Figures 3A and 3B are intended to be read as one figure when pieced together showing a flow diagram of one example of a method as employed in the invention.
Figure 4 is a more detailed schematic diagram of the bridge section of an accelerometer as employed in one embodiment of the invention.
Figure 5 is a more detailed schematic diagram of a portion of the sampling devices employed in one embodiment of the invention.
DESCRIPTION OF THE PREFERRED EMBODIMENT
Referring now to Figure 1, a block diagram of a cardiac pacemaker as provided by one embodiment of the invention is shown.
While it is deemed helpful to the understanding of the invention to describe it in terms of an application in a cardiac pacemaker, it is not intended to be so limited. Those skilled in the art will recognize other applications as uses for the invention in areas where low noise signal conditioning of low level, low frequency signals is advantageous. The pacemaker 1 includes a sensor 12, an analog signal processing section 2, a first means for summing 102, a second means for summing 104, a microcode processing means 110, a memory means 111, a delay means 120, a means for counting 130, a first clock 132, logical ANDgate means 140, and a means for generating stimulation pulses 150. The analog section 2 is coupled g 2~~~~~~
to the first summing means 102 by means of lines 303 and 301.
First and second summing means 102 and 104 share counting signals and are connected to microcode processor by means of SUMO signal line 307, SUM signal line 309, RESET signal line 311, and STOP
signal line 313. The pulse generator 150 is coupled to the microcode processor 110 by means of line 329 and branch line 315.
Line 329 is also connected to the delay means 120 by line 317. The delay means 120 is connected to the latching input of OUTI counter 130 by line 319. Clock 132 is connected to the clock input of OUTI
counter 130 by line 133. Memory means 111 is coupled by bus 113 to the microcode processor 110. Microcode processor 110 supplies data to OUTI counter 130 by means of bus 321. ANDgate 140 receives an output signal from the counter on line 325 and an ACTIVE signal from the microcode processor on line 323 and outputs a logical AND
function of those two signals to pulse generator 150 on line 327.
The memory means 111 is advantageously used to store parameters which may be preset by a physician and may, as a result, vary from patient to patient. Such parameters include a response factor (RESP) which, in one example embodiment of the invention, varies from .5 to 4 in 16 logarithmic steps.
The RESP factor is available to allow a physician to compensate for individual patients in which a pacemaker is implanted. For example, in an older individual, a small movement may require a greater increase in heart rate than the same movement in an athlete. However, a small movement will generate a small value for the sensor rectified average (SRA) because an older individual will not tend to move very rapidly. Since a large increase in heart rate is required for an older patient, a larger factor will be used to multiply up the change in rate for a given 3U SRA.
A base rate representing a minimum rate limit is also programmable by the physician. The physician may, in one example, select one of 16 preprogrammed rates as a base rate, where the 16 preprogrammed rates vary from 30 to 120 beats per minute. A
maximum sensor rate (MSR) may also be stored in memory to serve as an upper limit on the pulse generator output.
Referring now to Figure 2 the analog section 2 of the pacemaker of the invention is shown in a more detailed block diagram. The analog signal processing section is connected to a means for sensing acceleration 12 having programmable DC excitation current source 10 connected to a first excitation input 13 and a means for level shifting 20 connected to a second input 15.
In one example embodiment of the invention, the programmable DC excitation current means advantageously operates in approximately 4.7 microamp increments in 4 steps. The level shifting means 20 advantageously provides about a .5 volt level shift to the sensor 12. The sensor means 12 may advantageously be an integrated circuit accelerometer bonded in a leadless chip carrier. The accelerometer may advantageously include a balanced bridge that changes resistance in proportion to acceleration in the t2g range, perpendicular to the bonded surface, at frequencies from DC to 350 Hz. The accelerometer bridge may include a source resistance of about 3.5K ohms to 6.5K ohms. Such accelerometers are known in the art and are commercially available.
The sensor 12 provides first and second output signals 17, 19 to a means for chopping and modulating 26. Chopper modulator means 26 has first and second clock inputs 27 and 25 connected to a modulator clock 24. In one embodiment of the invention, the modulator clock operates at about 2.73 KHz. A first amplifier means 28 has first and second inputs 5 and 7 connected to first and 2~~~~~~
second outputs of the shopper modulator means 26 and further has an output 29 oonneotcd to a means far filtering 30. In one example embodiment, the first amplifier means 28 is a low »oise preamp halting t~ gaim of about 7. The filtering means 3o ia~ advantageously a band pass filter. Xn one example, the filteria~g means 30 advantageously comprises a fou>"th order filter herring a gain of about 9 and a frequency band of about 3o0 I~z to about $000 Hz. The level shifting means 20 is further connected by means of line 22 to a third input 22A of the amplifier meanr> ze and a level shifting input 22H of the band pass filte7cing maans~ 30. The band pass filtering mei~ns sled has a second input 4H and sin output 92. A
secotad amplifier 32 is connected to the output of the band pass Filter means 30 by means of a knee 3~.. an one example embodiment of-the invention, the second amplifier means has a gain of about 5 and compriass a wide swing amplifier. The second amplifier means has an output signal 33 which is ooup~.gd to a moans for demodulating 36. Demodulating means 36 includes a clock input 39 connected to a first filter clock 34. The first filter clock advantageously operates at about 2.~3 lct~z. cia:cuit elements 24, 28~ 30 and 36 operate to condition and amplify the sensor signal..
Ttaa demodulator output a5 is coupled to a low pass filter auEsaris 40 which also has a olack input 4~. connected to the filter sleek 34. The luw pass filter means 40 has an output 3~ which. icx connected by mearig of a lane 45 to means for attenuation 42.
Attsnuatiola means 42 includes an output 47 connected to a means for p~gset ~'emoval integration means 44. The offset removal intQgrat3.an means incsludes an output canneated by a line 99 to a modulation means 46. The attenuation means 42 and integrator means 44 each have olock inputs 51, 53 connected to a aeaond~ filter cloak 3(y 50. Tn ons example embodiment, the second Filter clock opsrates at ..
about 683 Hz. Modulation means 46 has a clock input 349 connected to modulator clock 24. The modulation means 46 further has an output 48 connected to the second input of band pass filter means 30.
The low pass filter means 40 is further connected by means of line 37A to a first input of a programmable gain amplifier means 56. The programmable gain amplifier means 56 further has a clock input at 43 which i;s also connected to the first filter clock 34.
Means for high pass filter integration 52 has an input 57 and an output 55 connected to the programmable gain amplifier at a second input. The programmable gain amplifier also includes an output 59 which is coupled by line 61 to input 57 of the high pass filter integrator. An analog-to-digital converter section 73 comprises a delta modulator 72, a zero crossing detector 74 and a digital output circuit 75. The delta modulator includes a clock input 71 connected to filter clock 34 and a first signal input 67. The zero crossing detector means 74 includes a clock input 73 connected to the filter clock 34 and a signal input 65: The delta modulator 72 outputs a signal on line 301 and the zero crossing detector 74 outputs a signal on line 303 to the digital output means 75.
Having described the configuration of the elements in the analog section, it will now prove helpful to the understanding of the invention to explain the operation of the analog section as implemented in ane example embodiment of the invention. It will be understood that this example is by way of explanation and not limitation of the invention. Still referring to Figure 2, in one example of the invention, the sensor means 12 is excited by the patient's motion which results in passing a current through the bridge section shown in detail in Figure 4 comprising resistors Rl, R2, R3 and R4. A differential acceleration signal is produced _ 7 _ across the bridge and is offset by a diode drop of about .5 volts by the level shifter 20 to move the signal into the operating range of the analog section 2. The signal at the sensor output is approximately 30 microvolts peak-to-peak/g riding on the 0.5 volt offset and is referred to as the Raw Accelerometer Output (RAO).
The expected voltage sensitivity at the bridge terminals 17 and 19 is about 30 microvolts per g of acceleration. The output signal transmitted to the chopper modulator 26 ranges in frequency from DC
up to about 350 Hz. The chopper modulator means 26 translates those frequencies up to about 2.73 ICHz by the modulation process using the modulator clock 24. Gain is added to the signal by amplifier means 28 which is preferably a low noise preamplifier with a gain of about 7. The band pass filter 30 then receives the amplified signal on line 29 and adds another gain of about 9. The second amplifier means 32 adds still another gain of about 9 to the signal before passing it to the demodulator 36. The band pass filter 30 is used for providing noise rejection of the chopper amplifier 26. This is required because the modulator 26 creates harmonics other than the chopper frequency which occur at integral multipliers of the chopper frequencies. The band pass filter 30 selects only those frequencies which have the best signal-to-noise ratio with frequencies in the range of about 300 to 8,000 Hertz.
After the initial modulation, amplification and filtering, the signal i.s transmitted on line 33 to the demodulator means 36. In one example embodiment, the signal is demodulated with a carrier of approximately 3 KHz. The demodulator 36 operates to demodulate the signal down to DC at output line 35. Any DC signal changes at the sensor are now present at the output of the demodulator. The demodulator does not operate to band limit the signal, therefore, the bandwidth is primarily limited by the sensor 12 up to this ., g _ point in the circuit. That is, it ranges in frequency from about 0 to about 300 Hz. On line 35, the signal has the same morphology as at the output of the sensor, however, the signal is increased in amplitude since it has been amplified. The amplifier stages initially receive about 1/3 of a microvolt input with reference to the RMS noise level. This allows the circuit to distinguish about 1/100 of a g and determine that it is something other than the system noise.
In ane example embodiment of the invention, the low pass filter 40 limits the highest frequencies of the system to about 10 Hz. At line 37, therefore, the signal frequency ranges from DC to 10 Hz. The low pass filter in one embodiment of the invention advantageously adds a gain of about 2 to the signal. The overall gain of the system is controlled by the programmable gain amplifier 56. In one embodiment of the invention, the programmable gain amplifier has 4 different settings. Relative settings of 1, 2, 4 or 8 may be selected through the gain select means 54. It is also possible to omit the programmable gain selection feature and, instead, fix the system gain to a pre-determined value.
The programmable gain amplifier means 56 and the high pass filter integration means 52 operate together to form a feedback loop which performs as a high pass filter. The high pass filter integration means 52 operates at a frequency of about 1.0 Hz. This feedback loop operates to eliminate DC signals. The bandwidth of the signal transmitted to the analog-to-digital converter 73 is thereby band limited to between 1 and 10 Hz. Attenuation means 42, offset removal integration means 44 and modulator 46 operate to provide a second feedback loop. This loop suppresses input signals that contain a large DC component which may be caused by DC errors in the accelerometer sensor 12 caused, in turn, by gravity or .. g .~
unbalanced bridges. The second feedback loop operates at a frequency of about .2 Hz. The second feedback loop feeds back the signal into a second input of the band pass filter having a gain of about .01. Essentially, the DC signal is reduced in amplitude and integrated through the second feedback loop. Because the band pass filter is operating at an RF frequency, that RF signal must be converted back up to the RF range with the modulator 46. After that, it is injected back into the band pass filter at line 48.
The DC removal loop comprising the attenuates 42, offset removal integrator 44 and modulator 46 operates at the front end of the analog section enabling suppression of DC signals before any amplifier reaches saturation. Therefore, the analog section can tolerate large DC components in the input signal at lines 17 and 19.
The second feedback loop comprising elements 42, 44 and 46 most particularly operates to remove instantaneous changes, as when the patient changes his body position in the earth field. For example, if the patient is lying on his back and rolls over to lie on his chest, the high pass filter 52 will remove the instantaneous change and the low pass filter 40 will recenter the operation of the front end amplifier 28 so that it is still capable of recognizing small signals. The sensor 12 is essentially DC and every different orientation in the field yields a different sensor output. Some such changes in orientation are usable signals.
The delta modulator means 72, zero crossing detector 74 and digital output means 75 operate together to form an analog-to-digital converter. The zero crossing detector 74 supplies signal rectification information by recognizing when the output signal crosses zero. At every zero crossing the delta modulator is purged of cumulative errors that may have been built up. Essentially, the l handoff between the analog and digital section comprises (1) zero crossings which are used to provide rectification and error removal in the delta modulator and (2) the delta modulator data stream.
The delta modulator 72 and zero crossing detector 74 may be constructed according to means well known in the art. The delta modulator provides fixed increments to the digital section. The delta modulator includes a tracking integrator which either increases or decreases by a fixed amount, based on a comparison between the integrator output and the input signal at line 67. For example, if the input at line 67 is greater than the integrator output at line 55, a "1" is produced in the delta modulator bit.
The logical "1" then increases the integrator value by one least significant bit (LSB). It keeps adding or subtracting one LSB or one count to track the input signal. Each time the signal increases, the delta modulator generates a "1" and each time the signal is decreasing the delta modulator generates a "0" . The bits are then counted over a predetermined interval to establish the value of the analog-to-digital converter. The second half of the analog-to-digital conversion occurs in the digital output section where the counting is accomplished.
The RAO is amplified by a programmable gain of 1000 - 8000 through the above-described circuitry and filtered by the programmable filter 56 with a nominal low pass of about 1 Hz and high pass at about 10 Hz. At output 59, the acceleration is approximately 600 Mv/g times the signal. This is referred to as the filtered accelerometer output (FAO). The FAO is digitized by means of the switched capacitor delta modulator circuit 72 sampling at a frequency of about 2738 Hz with a slew rate of.about 9 mv.
The deltas are rectified and counted by a 6 bit up/down counter at the same frequency, counter 102 yielding a range of 0-576 Mv.
~o~~~~~
Because of the rectification, only 0-300 My of the range is used.
The counter is constructed so that it will not roll over or under.
In addition, the up/down counter is reset when FAO crosses zero to keep the delta modulator from counting to the maximum or minimum value because of normal digitizing errors and offsets. The 6 bit value is passed to the SUM counter and is referred to as the Delta Modulator Output (DMO) at line 301.
Referring now to Figure 5, a more detailed block diagram of the digital output section 75 and summing sections 102, 104 of the digital section of the pacemaker are shown. An exclusive ORgate 402, a one-shot 406, and an up/down counter 404 comprise the digital output section 75. An adder 410 and a SUM latch 420 comprise the first counting means 102. The second counting means 104 is comprised of ripple counter 104. Delta modulator 72 provides a delta modulator signal on line 301 to one input of the exclusive ORgate 402. Zero crossing detector 74 provides a zero crossing signal on line 303 to a second input 403 of exclusive ORgate 402. Zero crossing detector also provides the zero crossing signal on input 401 of one shot 406 which is used to reset the up/down counter 404 by dividing the pulse on line 407. The up/down counter is operated at clock input 409 at a rate of about 2.7 KHz.
The exclusive ORgate provides a counting signal to the up/down counting input on line 405. The counter provides the delta modulator output on output 411 to the A input of adder 410. The output of adder 410 is routed on line 313 to the micrvcode processor 110 and on line 415 to the B input of adder 410. SUM
latch 420 continues to accumulate the 5UM until a reset pulse is received at line 437 from the microcode processor. Similarly, the SUMC counter is reset when a reset pulse from the rnicrocode processor is received on the reset input on line 439. Both the SUM
w accumulator and the SUMO counter are advantageously operated at about an 85 Hz rate which is provided on line 431 and routed to SUMC on line 433 and SUM 102 on lines 435. In one example embodiment of the invention, the SUMC counter is advantageously an 8 bit ripple counter. A sample clock 439 controls the sampling rate of the SUM and SUMC counters. In one example embodiment, the SUM
accumulator samples DMO every 11.72 msec (85 Hz) and adds the new DMO value to itself according to the equation: SUM = SUM + DMO.
The SUM accumulator will continue to add the DMO values to itself during each cardiac cycle. SUMC 104, counts the number of samples SUM during the cardiac cycle. At the end of the cardiac cycle, SUM
will range from 0 to 8192 (13 bits) and SUMC will range from 30 to 170 (8 bits)'. Every cardiac cycle, the microcode reads SUM and divides by SUMC according to the equation:
SRA = SUM -~ SUMC.
The final result is called the Sensor Rectified Average (SRA).
Because of the rectification described hereinabove and the energy function to be described below, only 10% of the range is used. The SUM is multiplied by 8 to move the SRA into a full dynamic range.
After SRA is formed, SUM and SUMO are reset and the summing and dividing cycle repeats. SRA ranges from 0, representing no activity, to 255 representing maximum activity. If SRA is calculated to be greater than 255, the microcode will force it to equal 255. A programmable option may advantageously be provided to even out the summing times between fast pacing rates and slow pacing rates. zf the pacing rate is faster than 512 msec., another cycle or two will be averaged before the microcode reads SUMO and SUM.
Because of noise or low level activity that may be sensed by the accelerometer, a Sensor Average Threshold (SAT) is advantageously included as programmable to ignore such unwanted values. This is done by subtracting SAT from SRA to give an Offset Sensor Rectified Average (OSRA). Once a number that represents the offset rectified average is produced, a Sensor Determined Interval (SDI) is formed using the Response Number (RESP) and subtracting it from the programmed rate parameter. OSRA is used with the programmable sensor response multiplier to form a sensor determined delta (SDD) which is rate in pulses per minute (PPM). The SDD is added to the programmed rate and the result is limited to the maximum sensor rate (MSR). After this step, the number is converted to the interval sensor determined interval (SDI) (where Interval = 60,000/ppm). SDI advantageously ranges from 255 (which corresponds to 30 ppm) to 45 (which corresponds to 170 ppm) and changes every cycle with activity. SDI is compared w~ah the actual pacing interval (API), to determine the response of the pulse generator. In simplified pairs, the algorithm can be thought of in three phases as follows: (1) if API is greater than SDI then the pacing interval is increased, but limited to SDI: (2) if API equals SDI then there is no change in the interval: (3) if API is less than SDI, then the interval is increased but limited to the sensor recovery number.
Now referring to Figures 3A and 38, a flow diagram of the miarocode contained in the mierocode processor is shown. The process begins at step 200 and continues to step 202 where the SUM
accumulator is stopped and read. At step 204 the value for SUM
read at step 202 is scaled. That is, since the value read is an average absolute value as opposed to a peak-to-peak value, it must be scaled upwards to optimize the dynamic range. At step 206, the SUM accumulator is reset and restarted. At step 208, the SRA value is determined by dividing SUM by the value in accumulator SUMC. At - 14 ..
step 210 the offset sensor rectified average OSRA is determined according to the method described hereinabove. The process then continues on to step 216 where a response factor (RESP) is applied to SBA to yield a change in rate called CRATE
At step 220, the target rate is calculated as the base rate value from a lookup table in memory added to ABATE. The target rate is limited to MSR by a comparison performed at step 224. The SDI value is then determined in milliseconds at step 226 as 60,000 divided by the limited target rate. The process continues to step 232 where SDI and API are compared. The first time through the algorithm an initial API value is set equal to the inverse of the base rate. If API is less than SDI, the process is routed to step 234. If API is greater than or equal to SDI, the process is routed to step 248. If it is not, the process proceeds to step 234 which executes a recovery algorithm in accordance with well known recovery methods. The process then continues to step 242 where API
is adjusted for 128 msec and output to the OUTI counter. Note that the 128 msec adjustment serves to account for the calculation time necessary to run the microcode algorithm. In effect, this adjustment removes a 128 msec offset from the start of a cardiac event caused by the time it takes to calculate the new API or actual pulsing interval. The 128 msec adjustment to API is advantageously accomplished by delay means 120. At step 244, telemetry and marker output selections are made in a conventional manner. At step 246 a cycle/SUM routine is entered and the process continues to step 300 to await the next standard logic chip reset pulse.
If step 248 is entered, a reaction algorithm sets an adjust value equal to API minus SDI over a preselected reaction factor.
The reaction factor is selected so as to avoid sudden large ~. ~~ c~
.~r ~ ... ~ :4 ~.
increases in heart rate. The process checks the adjust factor at step 250 and if it is less than 8 msec it is forced to 8 msec. API
is then set equal to API minus the adjust value but limited to SDI
at step 252. The process then proceeds as before to step 242.
Referring again to Figure 1, the adjusted and limited API
value is clocked into the OUTI counter 130 on bus 321. When the ACTIVE line 323 is high, the adjusted and limited AFI value is transmitted through gate l40 into the pulse generator which outputs stimulating pulses to electrodes (not shown) attached to the patient. The stimulating pulses so generated are at the rate inversely proportional to the adjusted API value.
This invention has been described herein in considerable detail in order to comply with the Patent Statutes and to provide those skilled in the art with the information needed to apply the novel principles and to construct and use such specialized components as are required. However, it is to be understood that the invention can be carried out by specifically different equipment and devices, and that various modifications, both as to the equipment details and operating procedures, can be accomplished without departing from the scope of the invention itself.
What is claimed is:
_ 1~ _
BACKGROUND OF THE INVENTION
I. Field of the Invention This invention relates to signal processing and rate modulation in automatic cardiac pacers and, more particularly, to a rate adaptive pacerwhich provides pacing signals to the heart based on a root mean square approximation of band limited, low frequency, low level acceleration signals generated from body motion.
II Description of the Prior Art Activity sensing rate-responsive pacemakers have been developed which use body vibration during physical activity as an indicator of the need to alter pacing rate. One such prior art device is described in US Patent No. 4,428,378 to Anderson, et al., entitled "Rate Adaptive Pacer". The Anderson, et al. pacemaker operates by responding to a variety of mechanical forces both internal and external to a pacemaker patient. Because the Anderson, et al. pacemaker responds to forces applied to the patient's body, it must necessarily rely on the outside environment to provide a mass to generate the measured forces. Such masses as provided by the outside environment are subject to wide variations. Yet another deficiency of such prior art sensors is that they can measure only changes in pressure, not constant pressure. The present invention provides an activity sensing rate-responsive pacemaker including an accelerometer which contains a captive, well defined, constant reference mass.
Prior art rate adaptive pacemakers which are dependent on mechanical force sensors have been found to be relatively insensitive to different levels of exertion, particularly with respect to providing appropriate rate changes in response to a patient ascending stairs and walking up inclines. Such p~~~cf ~~C~r -~-are also susceptible to the effects of extraneous vibration such as that encountered in various~forms of transport. One such study is reported by C.P. Lau, et al. in an article entitled "Selective Vibration Sensing: A New Concept for Activity - Sensing Rate-Responsive Pacing," P. 1299, PACE. Vol. II, Sept. 1988.
BUMMARY OF THE INVENTION
An accelerometer based rate adaptive pacemaker is provided.
Means for sensing acceleration is included wherein the acceleration sensing means provides an acceleration output signal corresponding to the sensed acceleration. Means for a sensor determined interval is coupled to the acceleration sensing means wherein the sensor determined interval is proportional to the acceleration output signal. Means for determining an actual pacing interval provides the actual pacing interval to a means for determining a pacing rate. The pacing rate means is, also, coupled to the sensor determined interval means. The pacing rate is a function of both the sensor determined interval and the actual pacing interval.
In one aspect of the invention, the programmed pacing interval is decreased if the actual pacing interval exceeds the sensor determined interval.
In another aspect of the invention, a root mean square (RMS) approximation is applied to a sensor output to yield a sensor rectified average value which is used for pacing rate modulation.
In yet another aspect of the invention, analog means are provided Eor amplifying and band limiting a low frequency, low level accelerometer signal with very low noise.
In still another aspect of the invention, a method is provided for converting signal energy into a pacing rate.
Other objects, features and advantages of the present invention will become apparent to those skilled in the art through the Description of the Preferred Embodiment, Claims, and drawings herein wherein like numerals refer to like elements.
BRIEF DESCRIPTION OF THE DRA'OPIN(38 Figure 1 is a block diagram of a cardiac pacemaker as provided by one embodiment of the invention.
Figure 2 is a block diagram of an analog signal processing network as employed in one embodiment of the invention.
Figures 3A and 3B are intended to be read as one figure when pieced together showing a flow diagram of one example of a method as employed in the invention.
Figure 4 is a more detailed schematic diagram of the bridge section of an accelerometer as employed in one embodiment of the invention.
Figure 5 is a more detailed schematic diagram of a portion of the sampling devices employed in one embodiment of the invention.
DESCRIPTION OF THE PREFERRED EMBODIMENT
Referring now to Figure 1, a block diagram of a cardiac pacemaker as provided by one embodiment of the invention is shown.
While it is deemed helpful to the understanding of the invention to describe it in terms of an application in a cardiac pacemaker, it is not intended to be so limited. Those skilled in the art will recognize other applications as uses for the invention in areas where low noise signal conditioning of low level, low frequency signals is advantageous. The pacemaker 1 includes a sensor 12, an analog signal processing section 2, a first means for summing 102, a second means for summing 104, a microcode processing means 110, a memory means 111, a delay means 120, a means for counting 130, a first clock 132, logical ANDgate means 140, and a means for generating stimulation pulses 150. The analog section 2 is coupled g 2~~~~~~
to the first summing means 102 by means of lines 303 and 301.
First and second summing means 102 and 104 share counting signals and are connected to microcode processor by means of SUMO signal line 307, SUM signal line 309, RESET signal line 311, and STOP
signal line 313. The pulse generator 150 is coupled to the microcode processor 110 by means of line 329 and branch line 315.
Line 329 is also connected to the delay means 120 by line 317. The delay means 120 is connected to the latching input of OUTI counter 130 by line 319. Clock 132 is connected to the clock input of OUTI
counter 130 by line 133. Memory means 111 is coupled by bus 113 to the microcode processor 110. Microcode processor 110 supplies data to OUTI counter 130 by means of bus 321. ANDgate 140 receives an output signal from the counter on line 325 and an ACTIVE signal from the microcode processor on line 323 and outputs a logical AND
function of those two signals to pulse generator 150 on line 327.
The memory means 111 is advantageously used to store parameters which may be preset by a physician and may, as a result, vary from patient to patient. Such parameters include a response factor (RESP) which, in one example embodiment of the invention, varies from .5 to 4 in 16 logarithmic steps.
The RESP factor is available to allow a physician to compensate for individual patients in which a pacemaker is implanted. For example, in an older individual, a small movement may require a greater increase in heart rate than the same movement in an athlete. However, a small movement will generate a small value for the sensor rectified average (SRA) because an older individual will not tend to move very rapidly. Since a large increase in heart rate is required for an older patient, a larger factor will be used to multiply up the change in rate for a given 3U SRA.
A base rate representing a minimum rate limit is also programmable by the physician. The physician may, in one example, select one of 16 preprogrammed rates as a base rate, where the 16 preprogrammed rates vary from 30 to 120 beats per minute. A
maximum sensor rate (MSR) may also be stored in memory to serve as an upper limit on the pulse generator output.
Referring now to Figure 2 the analog section 2 of the pacemaker of the invention is shown in a more detailed block diagram. The analog signal processing section is connected to a means for sensing acceleration 12 having programmable DC excitation current source 10 connected to a first excitation input 13 and a means for level shifting 20 connected to a second input 15.
In one example embodiment of the invention, the programmable DC excitation current means advantageously operates in approximately 4.7 microamp increments in 4 steps. The level shifting means 20 advantageously provides about a .5 volt level shift to the sensor 12. The sensor means 12 may advantageously be an integrated circuit accelerometer bonded in a leadless chip carrier. The accelerometer may advantageously include a balanced bridge that changes resistance in proportion to acceleration in the t2g range, perpendicular to the bonded surface, at frequencies from DC to 350 Hz. The accelerometer bridge may include a source resistance of about 3.5K ohms to 6.5K ohms. Such accelerometers are known in the art and are commercially available.
The sensor 12 provides first and second output signals 17, 19 to a means for chopping and modulating 26. Chopper modulator means 26 has first and second clock inputs 27 and 25 connected to a modulator clock 24. In one embodiment of the invention, the modulator clock operates at about 2.73 KHz. A first amplifier means 28 has first and second inputs 5 and 7 connected to first and 2~~~~~~
second outputs of the shopper modulator means 26 and further has an output 29 oonneotcd to a means far filtering 30. In one example embodiment, the first amplifier means 28 is a low »oise preamp halting t~ gaim of about 7. The filtering means 3o ia~ advantageously a band pass filter. Xn one example, the filteria~g means 30 advantageously comprises a fou>"th order filter herring a gain of about 9 and a frequency band of about 3o0 I~z to about $000 Hz. The level shifting means 20 is further connected by means of line 22 to a third input 22A of the amplifier meanr> ze and a level shifting input 22H of the band pass filte7cing maans~ 30. The band pass filtering mei~ns sled has a second input 4H and sin output 92. A
secotad amplifier 32 is connected to the output of the band pass Filter means 30 by means of a knee 3~.. an one example embodiment of-the invention, the second amplifier means has a gain of about 5 and compriass a wide swing amplifier. The second amplifier means has an output signal 33 which is ooup~.gd to a moans for demodulating 36. Demodulating means 36 includes a clock input 39 connected to a first filter clock 34. The first filter clock advantageously operates at about 2.~3 lct~z. cia:cuit elements 24, 28~ 30 and 36 operate to condition and amplify the sensor signal..
Ttaa demodulator output a5 is coupled to a low pass filter auEsaris 40 which also has a olack input 4~. connected to the filter sleek 34. The luw pass filter means 40 has an output 3~ which. icx connected by mearig of a lane 45 to means for attenuation 42.
Attsnuatiola means 42 includes an output 47 connected to a means for p~gset ~'emoval integration means 44. The offset removal intQgrat3.an means incsludes an output canneated by a line 99 to a modulation means 46. The attenuation means 42 and integrator means 44 each have olock inputs 51, 53 connected to a aeaond~ filter cloak 3(y 50. Tn ons example embodiment, the second Filter clock opsrates at ..
about 683 Hz. Modulation means 46 has a clock input 349 connected to modulator clock 24. The modulation means 46 further has an output 48 connected to the second input of band pass filter means 30.
The low pass filter means 40 is further connected by means of line 37A to a first input of a programmable gain amplifier means 56. The programmable gain amplifier means 56 further has a clock input at 43 which i;s also connected to the first filter clock 34.
Means for high pass filter integration 52 has an input 57 and an output 55 connected to the programmable gain amplifier at a second input. The programmable gain amplifier also includes an output 59 which is coupled by line 61 to input 57 of the high pass filter integrator. An analog-to-digital converter section 73 comprises a delta modulator 72, a zero crossing detector 74 and a digital output circuit 75. The delta modulator includes a clock input 71 connected to filter clock 34 and a first signal input 67. The zero crossing detector means 74 includes a clock input 73 connected to the filter clock 34 and a signal input 65: The delta modulator 72 outputs a signal on line 301 and the zero crossing detector 74 outputs a signal on line 303 to the digital output means 75.
Having described the configuration of the elements in the analog section, it will now prove helpful to the understanding of the invention to explain the operation of the analog section as implemented in ane example embodiment of the invention. It will be understood that this example is by way of explanation and not limitation of the invention. Still referring to Figure 2, in one example of the invention, the sensor means 12 is excited by the patient's motion which results in passing a current through the bridge section shown in detail in Figure 4 comprising resistors Rl, R2, R3 and R4. A differential acceleration signal is produced _ 7 _ across the bridge and is offset by a diode drop of about .5 volts by the level shifter 20 to move the signal into the operating range of the analog section 2. The signal at the sensor output is approximately 30 microvolts peak-to-peak/g riding on the 0.5 volt offset and is referred to as the Raw Accelerometer Output (RAO).
The expected voltage sensitivity at the bridge terminals 17 and 19 is about 30 microvolts per g of acceleration. The output signal transmitted to the chopper modulator 26 ranges in frequency from DC
up to about 350 Hz. The chopper modulator means 26 translates those frequencies up to about 2.73 ICHz by the modulation process using the modulator clock 24. Gain is added to the signal by amplifier means 28 which is preferably a low noise preamplifier with a gain of about 7. The band pass filter 30 then receives the amplified signal on line 29 and adds another gain of about 9. The second amplifier means 32 adds still another gain of about 9 to the signal before passing it to the demodulator 36. The band pass filter 30 is used for providing noise rejection of the chopper amplifier 26. This is required because the modulator 26 creates harmonics other than the chopper frequency which occur at integral multipliers of the chopper frequencies. The band pass filter 30 selects only those frequencies which have the best signal-to-noise ratio with frequencies in the range of about 300 to 8,000 Hertz.
After the initial modulation, amplification and filtering, the signal i.s transmitted on line 33 to the demodulator means 36. In one example embodiment, the signal is demodulated with a carrier of approximately 3 KHz. The demodulator 36 operates to demodulate the signal down to DC at output line 35. Any DC signal changes at the sensor are now present at the output of the demodulator. The demodulator does not operate to band limit the signal, therefore, the bandwidth is primarily limited by the sensor 12 up to this ., g _ point in the circuit. That is, it ranges in frequency from about 0 to about 300 Hz. On line 35, the signal has the same morphology as at the output of the sensor, however, the signal is increased in amplitude since it has been amplified. The amplifier stages initially receive about 1/3 of a microvolt input with reference to the RMS noise level. This allows the circuit to distinguish about 1/100 of a g and determine that it is something other than the system noise.
In ane example embodiment of the invention, the low pass filter 40 limits the highest frequencies of the system to about 10 Hz. At line 37, therefore, the signal frequency ranges from DC to 10 Hz. The low pass filter in one embodiment of the invention advantageously adds a gain of about 2 to the signal. The overall gain of the system is controlled by the programmable gain amplifier 56. In one embodiment of the invention, the programmable gain amplifier has 4 different settings. Relative settings of 1, 2, 4 or 8 may be selected through the gain select means 54. It is also possible to omit the programmable gain selection feature and, instead, fix the system gain to a pre-determined value.
The programmable gain amplifier means 56 and the high pass filter integration means 52 operate together to form a feedback loop which performs as a high pass filter. The high pass filter integration means 52 operates at a frequency of about 1.0 Hz. This feedback loop operates to eliminate DC signals. The bandwidth of the signal transmitted to the analog-to-digital converter 73 is thereby band limited to between 1 and 10 Hz. Attenuation means 42, offset removal integration means 44 and modulator 46 operate to provide a second feedback loop. This loop suppresses input signals that contain a large DC component which may be caused by DC errors in the accelerometer sensor 12 caused, in turn, by gravity or .. g .~
unbalanced bridges. The second feedback loop operates at a frequency of about .2 Hz. The second feedback loop feeds back the signal into a second input of the band pass filter having a gain of about .01. Essentially, the DC signal is reduced in amplitude and integrated through the second feedback loop. Because the band pass filter is operating at an RF frequency, that RF signal must be converted back up to the RF range with the modulator 46. After that, it is injected back into the band pass filter at line 48.
The DC removal loop comprising the attenuates 42, offset removal integrator 44 and modulator 46 operates at the front end of the analog section enabling suppression of DC signals before any amplifier reaches saturation. Therefore, the analog section can tolerate large DC components in the input signal at lines 17 and 19.
The second feedback loop comprising elements 42, 44 and 46 most particularly operates to remove instantaneous changes, as when the patient changes his body position in the earth field. For example, if the patient is lying on his back and rolls over to lie on his chest, the high pass filter 52 will remove the instantaneous change and the low pass filter 40 will recenter the operation of the front end amplifier 28 so that it is still capable of recognizing small signals. The sensor 12 is essentially DC and every different orientation in the field yields a different sensor output. Some such changes in orientation are usable signals.
The delta modulator means 72, zero crossing detector 74 and digital output means 75 operate together to form an analog-to-digital converter. The zero crossing detector 74 supplies signal rectification information by recognizing when the output signal crosses zero. At every zero crossing the delta modulator is purged of cumulative errors that may have been built up. Essentially, the l handoff between the analog and digital section comprises (1) zero crossings which are used to provide rectification and error removal in the delta modulator and (2) the delta modulator data stream.
The delta modulator 72 and zero crossing detector 74 may be constructed according to means well known in the art. The delta modulator provides fixed increments to the digital section. The delta modulator includes a tracking integrator which either increases or decreases by a fixed amount, based on a comparison between the integrator output and the input signal at line 67. For example, if the input at line 67 is greater than the integrator output at line 55, a "1" is produced in the delta modulator bit.
The logical "1" then increases the integrator value by one least significant bit (LSB). It keeps adding or subtracting one LSB or one count to track the input signal. Each time the signal increases, the delta modulator generates a "1" and each time the signal is decreasing the delta modulator generates a "0" . The bits are then counted over a predetermined interval to establish the value of the analog-to-digital converter. The second half of the analog-to-digital conversion occurs in the digital output section where the counting is accomplished.
The RAO is amplified by a programmable gain of 1000 - 8000 through the above-described circuitry and filtered by the programmable filter 56 with a nominal low pass of about 1 Hz and high pass at about 10 Hz. At output 59, the acceleration is approximately 600 Mv/g times the signal. This is referred to as the filtered accelerometer output (FAO). The FAO is digitized by means of the switched capacitor delta modulator circuit 72 sampling at a frequency of about 2738 Hz with a slew rate of.about 9 mv.
The deltas are rectified and counted by a 6 bit up/down counter at the same frequency, counter 102 yielding a range of 0-576 Mv.
~o~~~~~
Because of the rectification, only 0-300 My of the range is used.
The counter is constructed so that it will not roll over or under.
In addition, the up/down counter is reset when FAO crosses zero to keep the delta modulator from counting to the maximum or minimum value because of normal digitizing errors and offsets. The 6 bit value is passed to the SUM counter and is referred to as the Delta Modulator Output (DMO) at line 301.
Referring now to Figure 5, a more detailed block diagram of the digital output section 75 and summing sections 102, 104 of the digital section of the pacemaker are shown. An exclusive ORgate 402, a one-shot 406, and an up/down counter 404 comprise the digital output section 75. An adder 410 and a SUM latch 420 comprise the first counting means 102. The second counting means 104 is comprised of ripple counter 104. Delta modulator 72 provides a delta modulator signal on line 301 to one input of the exclusive ORgate 402. Zero crossing detector 74 provides a zero crossing signal on line 303 to a second input 403 of exclusive ORgate 402. Zero crossing detector also provides the zero crossing signal on input 401 of one shot 406 which is used to reset the up/down counter 404 by dividing the pulse on line 407. The up/down counter is operated at clock input 409 at a rate of about 2.7 KHz.
The exclusive ORgate provides a counting signal to the up/down counting input on line 405. The counter provides the delta modulator output on output 411 to the A input of adder 410. The output of adder 410 is routed on line 313 to the micrvcode processor 110 and on line 415 to the B input of adder 410. SUM
latch 420 continues to accumulate the 5UM until a reset pulse is received at line 437 from the microcode processor. Similarly, the SUMC counter is reset when a reset pulse from the rnicrocode processor is received on the reset input on line 439. Both the SUM
w accumulator and the SUMO counter are advantageously operated at about an 85 Hz rate which is provided on line 431 and routed to SUMC on line 433 and SUM 102 on lines 435. In one example embodiment of the invention, the SUMC counter is advantageously an 8 bit ripple counter. A sample clock 439 controls the sampling rate of the SUM and SUMC counters. In one example embodiment, the SUM
accumulator samples DMO every 11.72 msec (85 Hz) and adds the new DMO value to itself according to the equation: SUM = SUM + DMO.
The SUM accumulator will continue to add the DMO values to itself during each cardiac cycle. SUMC 104, counts the number of samples SUM during the cardiac cycle. At the end of the cardiac cycle, SUM
will range from 0 to 8192 (13 bits) and SUMC will range from 30 to 170 (8 bits)'. Every cardiac cycle, the microcode reads SUM and divides by SUMC according to the equation:
SRA = SUM -~ SUMC.
The final result is called the Sensor Rectified Average (SRA).
Because of the rectification described hereinabove and the energy function to be described below, only 10% of the range is used. The SUM is multiplied by 8 to move the SRA into a full dynamic range.
After SRA is formed, SUM and SUMO are reset and the summing and dividing cycle repeats. SRA ranges from 0, representing no activity, to 255 representing maximum activity. If SRA is calculated to be greater than 255, the microcode will force it to equal 255. A programmable option may advantageously be provided to even out the summing times between fast pacing rates and slow pacing rates. zf the pacing rate is faster than 512 msec., another cycle or two will be averaged before the microcode reads SUMO and SUM.
Because of noise or low level activity that may be sensed by the accelerometer, a Sensor Average Threshold (SAT) is advantageously included as programmable to ignore such unwanted values. This is done by subtracting SAT from SRA to give an Offset Sensor Rectified Average (OSRA). Once a number that represents the offset rectified average is produced, a Sensor Determined Interval (SDI) is formed using the Response Number (RESP) and subtracting it from the programmed rate parameter. OSRA is used with the programmable sensor response multiplier to form a sensor determined delta (SDD) which is rate in pulses per minute (PPM). The SDD is added to the programmed rate and the result is limited to the maximum sensor rate (MSR). After this step, the number is converted to the interval sensor determined interval (SDI) (where Interval = 60,000/ppm). SDI advantageously ranges from 255 (which corresponds to 30 ppm) to 45 (which corresponds to 170 ppm) and changes every cycle with activity. SDI is compared w~ah the actual pacing interval (API), to determine the response of the pulse generator. In simplified pairs, the algorithm can be thought of in three phases as follows: (1) if API is greater than SDI then the pacing interval is increased, but limited to SDI: (2) if API equals SDI then there is no change in the interval: (3) if API is less than SDI, then the interval is increased but limited to the sensor recovery number.
Now referring to Figures 3A and 38, a flow diagram of the miarocode contained in the mierocode processor is shown. The process begins at step 200 and continues to step 202 where the SUM
accumulator is stopped and read. At step 204 the value for SUM
read at step 202 is scaled. That is, since the value read is an average absolute value as opposed to a peak-to-peak value, it must be scaled upwards to optimize the dynamic range. At step 206, the SUM accumulator is reset and restarted. At step 208, the SRA value is determined by dividing SUM by the value in accumulator SUMC. At - 14 ..
step 210 the offset sensor rectified average OSRA is determined according to the method described hereinabove. The process then continues on to step 216 where a response factor (RESP) is applied to SBA to yield a change in rate called CRATE
At step 220, the target rate is calculated as the base rate value from a lookup table in memory added to ABATE. The target rate is limited to MSR by a comparison performed at step 224. The SDI value is then determined in milliseconds at step 226 as 60,000 divided by the limited target rate. The process continues to step 232 where SDI and API are compared. The first time through the algorithm an initial API value is set equal to the inverse of the base rate. If API is less than SDI, the process is routed to step 234. If API is greater than or equal to SDI, the process is routed to step 248. If it is not, the process proceeds to step 234 which executes a recovery algorithm in accordance with well known recovery methods. The process then continues to step 242 where API
is adjusted for 128 msec and output to the OUTI counter. Note that the 128 msec adjustment serves to account for the calculation time necessary to run the microcode algorithm. In effect, this adjustment removes a 128 msec offset from the start of a cardiac event caused by the time it takes to calculate the new API or actual pulsing interval. The 128 msec adjustment to API is advantageously accomplished by delay means 120. At step 244, telemetry and marker output selections are made in a conventional manner. At step 246 a cycle/SUM routine is entered and the process continues to step 300 to await the next standard logic chip reset pulse.
If step 248 is entered, a reaction algorithm sets an adjust value equal to API minus SDI over a preselected reaction factor.
The reaction factor is selected so as to avoid sudden large ~. ~~ c~
.~r ~ ... ~ :4 ~.
increases in heart rate. The process checks the adjust factor at step 250 and if it is less than 8 msec it is forced to 8 msec. API
is then set equal to API minus the adjust value but limited to SDI
at step 252. The process then proceeds as before to step 242.
Referring again to Figure 1, the adjusted and limited API
value is clocked into the OUTI counter 130 on bus 321. When the ACTIVE line 323 is high, the adjusted and limited AFI value is transmitted through gate l40 into the pulse generator which outputs stimulating pulses to electrodes (not shown) attached to the patient. The stimulating pulses so generated are at the rate inversely proportional to the adjusted API value.
This invention has been described herein in considerable detail in order to comply with the Patent Statutes and to provide those skilled in the art with the information needed to apply the novel principles and to construct and use such specialized components as are required. However, it is to be understood that the invention can be carried out by specifically different equipment and devices, and that various modifications, both as to the equipment details and operating procedures, can be accomplished without departing from the scope of the invention itself.
What is claimed is:
_ 1~ _
Claims (15)
1. An accelerometer based rate adaptive pacemaker comprising:
(a) means for sensing acceleration wherein the acceleration sensing means provides an acceleration output signal corresponding to the sensed acceleration;
(b) means for periodically sampling the acceleration output signal and for summing the samples over a cardiac cycle to obtain an accumulated sum;
(c) means for computing a sensor rectified average as the quotient of the accumulated sum divided by the number of samples acquired in the cardiac cycle;
(d) means including the means for computing for providing a sensor determined interval from the sensor rectified average;
(e) means for determining an actual pacing interval; and (f) means for determining a pacing rate coupled to the means for providing the sensor determined interval and the means for determining the actual pacing interval, wherein the pacing rate is a function of both the sensor determined interval and the actual pacing interval.
(a) means for sensing acceleration wherein the acceleration sensing means provides an acceleration output signal corresponding to the sensed acceleration;
(b) means for periodically sampling the acceleration output signal and for summing the samples over a cardiac cycle to obtain an accumulated sum;
(c) means for computing a sensor rectified average as the quotient of the accumulated sum divided by the number of samples acquired in the cardiac cycle;
(d) means including the means for computing for providing a sensor determined interval from the sensor rectified average;
(e) means for determining an actual pacing interval; and (f) means for determining a pacing rate coupled to the means for providing the sensor determined interval and the means for determining the actual pacing interval, wherein the pacing rate is a function of both the sensor determined interval and the actual pacing interval.
2. The pacemaker of Claim 1 wherein the pacing rate is increased if the actual pacing interval exceeds the sensor determined interval.
3. The pacemaker of Claim 2 wherein the pacing rate is maintained at its current value if the actual pacing interval is equal to the sensor determined interval.
4. The pacemaker of Claim 3 wherein the pacing rate is decreased if the actual pacing interval is less than the sensor pacing interval.
5. The pacemaker of Claim 1 wherein the means for sensing acceleration comprises an integrated accelerometer.
6. The pacemaker of Claim 5 wherein the means for providing a sensor determined interval comprises an analog section and a digital section.
7. The pacemaker of Claim 6 wherein the analog section comprises:
(a) means for chopping and modulating the acceleration output signal and providing a first modulated signal, wherein the chopping and modulating means is connected to receive the acceleration output signal;
(b) first means for amplifying coupled to the chopping and modulating means to receive the first modulated signal and provide a first amplified output;
(c) first means for filtering having a first input coupled to receive the first amplified output, having a second input and including a first filtered output;
(d) second means for amplifying coupled to the first filtered output and providing a second amplified output;
(e) demodulator means for demodulating coupled at an input to the second amplified output and providing a demodulated output;
(f) second means for filtering having an input coupled to the demodulated output and providing a second filtered output;
(g) feedback means for suppressing DC signals having an input coupled to the second filtered output and having an output coupled to the second input of the first filtering means;
(h) third means for amplifying also coupled to the second filtered output at a first input including a second input and a third amplified output;
(i) means for filtering and integrating having an input coupled to the third amplifier output and having an output coupled to the second input of the third amplifier; and (j) means for converting analog signals to digital signals having first and second inputs coupled to the third amplified output.
(a) means for chopping and modulating the acceleration output signal and providing a first modulated signal, wherein the chopping and modulating means is connected to receive the acceleration output signal;
(b) first means for amplifying coupled to the chopping and modulating means to receive the first modulated signal and provide a first amplified output;
(c) first means for filtering having a first input coupled to receive the first amplified output, having a second input and including a first filtered output;
(d) second means for amplifying coupled to the first filtered output and providing a second amplified output;
(e) demodulator means for demodulating coupled at an input to the second amplified output and providing a demodulated output;
(f) second means for filtering having an input coupled to the demodulated output and providing a second filtered output;
(g) feedback means for suppressing DC signals having an input coupled to the second filtered output and having an output coupled to the second input of the first filtering means;
(h) third means for amplifying also coupled to the second filtered output at a first input including a second input and a third amplified output;
(i) means for filtering and integrating having an input coupled to the third amplifier output and having an output coupled to the second input of the third amplifier; and (j) means for converting analog signals to digital signals having first and second inputs coupled to the third amplified output.
8. The pacemaker of Claim 7 wherein the feedback means comprises:
(a) means for attenuation having an input coupled to the second filtered and having an attenuated output;
(b) means for removing offset signals having an input coupled to the attenuated output and having an offset removal output; and (c) means for modulating having an input coupled to the offset removal output and having an output coupled to the second input of the first filtering means.
(a) means for attenuation having an input coupled to the second filtered and having an attenuated output;
(b) means for removing offset signals having an input coupled to the attenuated output and having an offset removal output; and (c) means for modulating having an input coupled to the offset removal output and having an output coupled to the second input of the first filtering means.
9. The pacemaker of Claim 7 wherein the third amplifying means comprises a programmable gain amplifier.
10. The pacemaker of Claim 7 wherein the first filtering means comprises a band pass filter having a bandwidth ranging from about 300 Hz to 8000 Hz.
11. The pacemaker of Claim 7 wherein the analog-to-digital conversion means comprises a delta modulator coupled to the first input and a zero crossing detector coupled at the second input.
12. The pacemaker of Claim 8 wherein the analog-to-digital conversion means comprises a delta modulator coupled to the first input and a zero crossing detector coupled at the second input.
13. A method for operating a cardiac pacemaker comprising the steps of:
(a) sensing acceleration of the pacemaker;
(b) providing an acceleration output signal proportional to the sensed acceleration;
(c) band limiting and filtering the acceleration output signal to provide a filtered accelerator output;
(d) digitizing the filtered accelerator output;
(e) sampling the filtered accelerator output and summing the samples over a cardiac cycle to provide an accumulated sum;
(f) determining a sensor rectified average proportional to the accumulated sum divided by the number of samples in a cardiac cycle;
(g) determining a sensor determined interval proportional to the sensor rectified average; and (h) determining a pacing rate based in part on the sensor determined interval.
(a) sensing acceleration of the pacemaker;
(b) providing an acceleration output signal proportional to the sensed acceleration;
(c) band limiting and filtering the acceleration output signal to provide a filtered accelerator output;
(d) digitizing the filtered accelerator output;
(e) sampling the filtered accelerator output and summing the samples over a cardiac cycle to provide an accumulated sum;
(f) determining a sensor rectified average proportional to the accumulated sum divided by the number of samples in a cardiac cycle;
(g) determining a sensor determined interval proportional to the sensor rectified average; and (h) determining a pacing rate based in part on the sensor determined interval.
14. The method of Claim 13 wherein the step for determining a pacing rate, further comprises the steps of:
(a) determining an actual pacing interval;
(b) comparing the actual pacing interval with the sensor determined interval;
(c) executing a recovery algorithm if the actual pacing interval is less than the sensor determined interval; and (d) executing a reaction algorithm if the actual pacing interval is greater than or equal to the sensor determined interval.
(a) determining an actual pacing interval;
(b) comparing the actual pacing interval with the sensor determined interval;
(c) executing a recovery algorithm if the actual pacing interval is less than the sensor determined interval; and (d) executing a reaction algorithm if the actual pacing interval is greater than or equal to the sensor determined interval.
15. The method of Claim 14 wherein the step of executing a reaction algorithm comprises the steps of:
(a) determining an adjustment factor proportional to the difference between the actual pacing interval and the sensor determined interval;
(b) adjusting the actual pacing interval by applying the adjustment factor to provide an adjusted actual pacing interval;
(c) compensating the adjusted actual pacing interval for processing time delays; and (d) outputting the adjusted and compensated actual pacing interval as the pacing rate to the cardiac pacemaker.
(a) determining an adjustment factor proportional to the difference between the actual pacing interval and the sensor determined interval;
(b) adjusting the actual pacing interval by applying the adjustment factor to provide an adjusted actual pacing interval;
(c) compensating the adjusted actual pacing interval for processing time delays; and (d) outputting the adjusted and compensated actual pacing interval as the pacing rate to the cardiac pacemaker.
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| US07/641,549 | 1991-01-15 | ||
| US07/641,549 US5179947A (en) | 1991-01-15 | 1991-01-15 | Acceleration-sensitive cardiac pacemaker and method of operation |
Publications (2)
| Publication Number | Publication Date |
|---|---|
| CA2054941A1 CA2054941A1 (en) | 1992-07-16 |
| CA2054941C true CA2054941C (en) | 2001-08-07 |
Family
ID=24572865
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA002054941A Expired - Fee Related CA2054941C (en) | 1991-01-15 | 1991-11-05 | Cardiac pacemaker |
Country Status (5)
| Country | Link |
|---|---|
| US (1) | US5179947A (en) |
| EP (1) | EP0495293B2 (en) |
| JP (1) | JP2535115B2 (en) |
| CA (1) | CA2054941C (en) |
| DE (1) | DE69122077T3 (en) |
Families Citing this family (80)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US5360436A (en) * | 1992-04-03 | 1994-11-01 | Intermedics, Inc. | Cardiac pacing responsive to multiple activity types |
| US5370667A (en) * | 1992-04-03 | 1994-12-06 | Intermedics, Inc. | Device and method for automatically adjusting tachycardia recognition criteria based on detected parameter |
| IT1256900B (en) * | 1992-07-27 | 1995-12-27 | Franco Vallana | PROCEDURE AND DEVICE TO DETECT CARDIAC FUNCTIONALITY. |
| US5413592A (en) * | 1993-03-26 | 1995-05-09 | Intermedics, Inc. | Cardiac pacemaker with automatic parameter adjustment |
| IT1260692B (en) * | 1993-10-05 | 1996-04-22 | Sorin Biomedica Spa | DEVICE FOR DETERMINING THE MYOCARDIC FUNCTION AND ITS PROCEDURE. |
| US5609612A (en) * | 1993-10-05 | 1997-03-11 | Sorin Biomedica Cardio S.P.A. | Device for determining myocardial function and corresponding procedure and method |
| US5649968A (en) * | 1995-11-14 | 1997-07-22 | Intermedics, Inc. | Accelerometer-based rate-adaptive cardiac pacing with second generation signal processing |
| US5978972A (en) * | 1996-06-14 | 1999-11-09 | Johns Hopkins University | Helmet system including at least three accelerometers and mass memory and method for recording in real-time orthogonal acceleration data of a head |
| US5792195A (en) * | 1996-12-16 | 1998-08-11 | Cardiac Pacemakers, Inc. | Acceleration sensed safe upper rate envelope for calculating the hemodynamic upper rate limit for a rate adaptive cardiac rhythm management device |
| US5792183A (en) * | 1997-01-27 | 1998-08-11 | Cardiac Pacemakers, Inc. | Combination pacemaker and defibrillator having dynamic ventricular refractory period |
| IL120507A (en) * | 1997-03-24 | 2001-06-14 | Keytron Electronics & Technolo | Exercise monitoring system |
| AU3732000A (en) * | 1999-03-12 | 2000-09-28 | Cardiac Pacemakers, Inc. | Cardiac rhythm management system with time-dependent frequency response |
| DE19929553A1 (en) * | 1999-06-23 | 2001-01-04 | Biotronik Mess & Therapieg | Pacemaker |
| US6317626B1 (en) | 1999-11-03 | 2001-11-13 | Medtronic, Inc. | Method and apparatus for monitoring heart rate |
| JP3977569B2 (en) * | 2000-03-06 | 2007-09-19 | テルモ株式会社 | Biological implantable electrode lead and biological implantable medical device using the same |
| US7191000B2 (en) * | 2001-07-31 | 2007-03-13 | Cardiac Pacemakers, Inc. | Cardiac rhythm management system for edema |
| US6723077B2 (en) * | 2001-09-28 | 2004-04-20 | Hewlett-Packard Development Company, L.P. | Cutaneous administration system |
| US6973349B2 (en) * | 2001-12-05 | 2005-12-06 | Cardiac Pacemakers, Inc. | Method and apparatus for minimizing post-infarct ventricular remodeling |
| US8391989B2 (en) | 2002-12-18 | 2013-03-05 | Cardiac Pacemakers, Inc. | Advanced patient management for defining, identifying and using predetermined health-related events |
| US7113825B2 (en) * | 2002-05-03 | 2006-09-26 | Cardiac Pacemakers, Inc. | Method and apparatus for detecting acoustic oscillations in cardiac rhythm |
| US7123962B2 (en) * | 2002-12-02 | 2006-10-17 | Cardiac Pacemakers, Inc. | Phonocardiographic image-based atrioventricular delay optimization |
| US7972275B2 (en) * | 2002-12-30 | 2011-07-05 | Cardiac Pacemakers, Inc. | Method and apparatus for monitoring of diastolic hemodynamics |
| US7248923B2 (en) * | 2003-11-06 | 2007-07-24 | Cardiac Pacemakers, Inc. | Dual-use sensor for rate responsive pacing and heart sound monitoring |
| US7215997B2 (en) * | 2003-12-22 | 2007-05-08 | Cardiac Pacemakers, Inc. | Dynamic device therapy control for treating post myocardial infarction patients |
| US7431699B2 (en) * | 2003-12-24 | 2008-10-07 | Cardiac Pacemakers, Inc. | Method and apparatus for third heart sound detection |
| US7115096B2 (en) | 2003-12-24 | 2006-10-03 | Cardiac Pacemakers, Inc. | Third heart sound activity index for heart failure monitoring |
| US7840263B2 (en) * | 2004-02-27 | 2010-11-23 | Cardiac Pacemakers, Inc. | Method and apparatus for device controlled gene expression |
| US7209786B2 (en) | 2004-06-10 | 2007-04-24 | Cardiac Pacemakers, Inc. | Method and apparatus for optimization of cardiac resynchronization therapy using heart sounds |
| US7480528B2 (en) * | 2004-07-23 | 2009-01-20 | Cardiac Pacemakers, Inc. | Method and apparatus for monitoring heart failure patients with cardiopulmonary comorbidities |
| US7559901B2 (en) | 2004-07-28 | 2009-07-14 | Cardiac Pacemakers, Inc. | Determining a patient's posture from mechanical vibrations of the heart |
| US7387610B2 (en) | 2004-08-19 | 2008-06-17 | Cardiac Pacemakers, Inc. | Thoracic impedance detection with blood resistivity compensation |
| US7662104B2 (en) | 2005-01-18 | 2010-02-16 | Cardiac Pacemakers, Inc. | Method for correction of posture dependence on heart sounds |
| US7603170B2 (en) * | 2005-04-26 | 2009-10-13 | Cardiac Pacemakers, Inc. | Calibration of impedance monitoring of respiratory volumes using thoracic D.C. impedance |
| US7907997B2 (en) * | 2005-05-11 | 2011-03-15 | Cardiac Pacemakers, Inc. | Enhancements to the detection of pulmonary edema when using transthoracic impedance |
| US7340296B2 (en) | 2005-05-18 | 2008-03-04 | Cardiac Pacemakers, Inc. | Detection of pleural effusion using transthoracic impedance |
| US7922669B2 (en) | 2005-06-08 | 2011-04-12 | Cardiac Pacemakers, Inc. | Ischemia detection using a heart sound sensor |
| EP2471452B1 (en) * | 2005-10-14 | 2014-12-10 | Pacesetter, Inc. | Cardiac pacing system and method of conveying information therein |
| US9168383B2 (en) | 2005-10-14 | 2015-10-27 | Pacesetter, Inc. | Leadless cardiac pacemaker with conducted communication |
| US8108034B2 (en) | 2005-11-28 | 2012-01-31 | Cardiac Pacemakers, Inc. | Systems and methods for valvular regurgitation detection |
| US7780606B2 (en) * | 2006-03-29 | 2010-08-24 | Cardiac Pacemakers, Inc. | Hemodynamic stability assessment based on heart sounds |
| US8000780B2 (en) * | 2006-06-27 | 2011-08-16 | Cardiac Pacemakers, Inc. | Detection of myocardial ischemia from the time sequence of implanted sensor measurements |
| US7532927B1 (en) * | 2006-07-19 | 2009-05-12 | Pacesetter, Inc. | Implantable medical device and associated method for reducing an affect of a vibratory alert on a vibration-sensitive sensor |
| US20080119749A1 (en) | 2006-11-20 | 2008-05-22 | Cardiac Pacemakers, Inc. | Respiration-synchronized heart sound trending |
| US8096954B2 (en) | 2006-11-29 | 2012-01-17 | Cardiac Pacemakers, Inc. | Adaptive sampling of heart sounds |
| US9615744B2 (en) | 2007-01-31 | 2017-04-11 | Medtronic, Inc. | Chopper-stabilized instrumentation amplifier for impedance measurement |
| US7385443B1 (en) * | 2007-01-31 | 2008-06-10 | Medtronic, Inc. | Chopper-stabilized instrumentation amplifier |
| US7391257B1 (en) | 2007-01-31 | 2008-06-24 | Medtronic, Inc. | Chopper-stabilized instrumentation amplifier for impedance measurement |
| US8265769B2 (en) * | 2007-01-31 | 2012-09-11 | Medtronic, Inc. | Chopper-stabilized instrumentation amplifier for wireless telemetry |
| US7853327B2 (en) | 2007-04-17 | 2010-12-14 | Cardiac Pacemakers, Inc. | Heart sound tracking system and method |
| US8781595B2 (en) | 2007-04-30 | 2014-07-15 | Medtronic, Inc. | Chopper mixer telemetry circuit |
| US8594779B2 (en) * | 2007-04-30 | 2013-11-26 | Medtronic, Inc. | Seizure prediction |
| US9788750B2 (en) * | 2007-04-30 | 2017-10-17 | Medtronic, Inc. | Seizure prediction |
| EP2203216A1 (en) * | 2007-09-20 | 2010-07-07 | Nanostim, Inc. | Leadless cardiac pacemaker with secondary fixation capability |
| US8380314B2 (en) | 2007-09-26 | 2013-02-19 | Medtronic, Inc. | Patient directed therapy control |
| WO2009042172A2 (en) * | 2007-09-26 | 2009-04-02 | Medtronic, Inc. | Frequency selective monitoring of physiological signals |
| US7714757B2 (en) * | 2007-09-26 | 2010-05-11 | Medtronic, Inc. | Chopper-stabilized analog-to-digital converter |
| US7623053B2 (en) * | 2007-09-26 | 2009-11-24 | Medtronic, Inc. | Implantable medical device with low power delta-sigma analog-to-digital converter |
| US8121694B2 (en) | 2007-10-16 | 2012-02-21 | Medtronic, Inc. | Therapy control based on a patient movement state |
| CN101925377A (en) | 2008-01-25 | 2010-12-22 | 麦德托尼克公司 | The detection of Sleep stages |
| US8478402B2 (en) * | 2008-10-31 | 2013-07-02 | Medtronic, Inc. | Determining intercardiac impedance |
| US20100113964A1 (en) * | 2008-10-31 | 2010-05-06 | Wahlstrand John D | Determining intercardiac impedance |
| WO2010088687A1 (en) | 2009-02-02 | 2010-08-05 | Nanostim, Inc. | Leadless cardiac pacemaker with secondary fixation capability |
| US20110077708A1 (en) * | 2009-09-28 | 2011-03-31 | Alan Ostroff | MRI Compatible Leadless Cardiac Pacemaker |
| US9770204B2 (en) | 2009-11-11 | 2017-09-26 | Medtronic, Inc. | Deep brain stimulation for sleep and movement disorders |
| US9060692B2 (en) | 2010-10-12 | 2015-06-23 | Pacesetter, Inc. | Temperature sensor for a leadless cardiac pacemaker |
| EP2627403A4 (en) | 2010-10-12 | 2014-03-26 | Nanostim Inc | Temperature sensor for a leadless cardiac pacemaker |
| WO2012051235A1 (en) | 2010-10-13 | 2012-04-19 | Nanostim, Inc. | Leadless cardiac pacemaker with anti-unscrewing feature |
| US9132275B2 (en) | 2010-11-18 | 2015-09-15 | Cardiac Pacemakers, Inc. | Automatic determination of chronotropic incompetence using atrial pacing at rest |
| EP2651502B1 (en) | 2010-12-13 | 2016-11-09 | Pacesetter, Inc. | Pacemaker retrieval systems |
| EP2651494B1 (en) | 2010-12-13 | 2017-02-15 | Pacesetter, Inc. | Delivery catheter |
| EP2654889B1 (en) | 2010-12-20 | 2017-03-01 | Pacesetter, Inc. | Leadless pacemaker with radial fixation mechanism |
| WO2013067496A2 (en) | 2011-11-04 | 2013-05-10 | Nanostim, Inc. | Leadless cardiac pacemaker with integral battery and redundant welds |
| CN102526879B (en) * | 2012-02-16 | 2014-08-06 | 上海微创医疗器械（集团）有限公司 | Device and system for adaptively adjusting pacing frequency |
| EP2879758B1 (en) | 2012-08-01 | 2018-04-18 | Pacesetter, Inc. | Biostimulator circuit with flying cell |
| US9439150B2 (en) | 2013-03-15 | 2016-09-06 | Medtronic, Inc. | Control of spectral agressors in a physiological signal montoring device |
| US9521979B2 (en) | 2013-03-15 | 2016-12-20 | Medtronic, Inc. | Control of spectral agressors in a physiological signal monitoring device |
| US9399140B2 (en) | 2014-07-25 | 2016-07-26 | Medtronic, Inc. | Atrial contraction detection by a ventricular leadless pacing device for atrio-synchronous ventricular pacing |
| US9924904B2 (en) | 2014-09-02 | 2018-03-27 | Medtronic, Inc. | Power-efficient chopper amplifier |
| DE102015109037A1 (en) | 2015-06-09 | 2016-12-15 | Biotronik Se & Co. Kg | Rate-adaptive intra- or epicardial cardiac stimulator and activity sensor |
| WO2018009569A1 (en) | 2016-07-06 | 2018-01-11 | Cardiac Pacemakers, Inc. | Method and system for determining an atrial contraction timing fiducial in a leadless cardiac pacemaker system |
Family Cites Families (14)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US4428378A (en) * | 1981-11-19 | 1984-01-31 | Medtronic, Inc. | Rate adaptive pacer |
| DE3226345A1 (en) * | 1982-07-14 | 1984-01-19 | Siemens AG, 1000 Berlin und 8000 München | Implantable pacemaker having a detector controlling the same |
| DE3506789A1 (en) * | 1985-02-22 | 1986-08-28 | Biotronik Meß- und Therapiegeräte GmbH & Co Ingenieurbüro Berlin, 1000 Berlin | DEMAND HEART PACEMAKER WITH PHYSIOLOGICAL CONTROL |
| US4702253A (en) * | 1985-10-15 | 1987-10-27 | Telectronics N.V. | Metabolic-demand pacemaker and method of using the same to determine minute volume |
| US4722342A (en) * | 1986-06-16 | 1988-02-02 | Siemens Aktiengesellschaft | Cardiac pacer for pacing a human heart and pacing method |
| DE3631155A1 (en) * | 1986-09-12 | 1988-03-24 | Alt Eckhard | FREQUENCY VARIABLE HEART PACEMAKER WITH STRESS-ADEQUATE FREQUENCY BEHAVIOR |
| US4886064A (en) * | 1987-11-25 | 1989-12-12 | Siemens Aktiengesellschaft | Body activity controlled heart pacer |
| DE68927447T2 (en) * | 1988-02-17 | 1997-04-03 | Stuart Charles Webb | Frequency-sensitive pacemaker |
| DE3822662A1 (en) * | 1988-07-05 | 1990-01-11 | Bosch Gmbh Robert | PNEUMATICALLY OPERATED PRESSURE SWITCH AND METHOD FOR THE PRODUCTION THEREOF |
| US4972834A (en) * | 1988-09-30 | 1990-11-27 | Vitatron Medical B.V. | Pacemaker with improved dynamic rate responsiveness |
| US4940053A (en) * | 1989-01-25 | 1990-07-10 | Siemens-Pacesetter, Inc. | Energy controlled rate-responsive pacemaker having automatically adjustable control parameters |
| US4940052A (en) * | 1989-01-25 | 1990-07-10 | Siemens-Pacesetter, Inc. | Microprocessor controlled rate-responsive pacemaker having automatic rate response threshold adjustment |
| US4922930A (en) * | 1989-04-11 | 1990-05-08 | Intermedics, Inc. | Implantable device with circadian rhythm adjustment |
| US5027813A (en) * | 1990-03-05 | 1991-07-02 | Cardiac Pacemakers, Inc. | Rate responsive pacemaker apparatus having an electrode interface sensor |
-
1991
- 1991-01-15 US US07/641,549 patent/US5179947A/en not_active Expired - Lifetime
- 1991-09-27 EP EP91308893A patent/EP0495293B2/en not_active Expired - Lifetime
- 1991-09-27 DE DE69122077T patent/DE69122077T3/en not_active Expired - Fee Related
- 1991-11-05 CA CA002054941A patent/CA2054941C/en not_active Expired - Fee Related
- 1991-12-17 JP JP3333692A patent/JP2535115B2/en not_active Expired - Lifetime
Also Published As
| Publication number | Publication date |
|---|---|
| JPH04303464A (en) | 1992-10-27 |
| DE69122077D1 (en) | 1996-10-17 |
| JP2535115B2 (en) | 1996-09-18 |
| DE69122077T3 (en) | 1999-12-02 |
| EP0495293B1 (en) | 1996-09-11 |
| CA2054941A1 (en) | 1992-07-16 |
| US5179947A (en) | 1993-01-19 |
| DE69122077T2 (en) | 1997-01-23 |
| EP0495293B2 (en) | 2000-07-05 |
| EP0495293A1 (en) | 1992-07-22 |
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