U.S. patent application number 12/318555 was filed with the patent office on 2009-04-30 for method and device for monitoring and improving patient-ventilator interaction.
Invention is credited to Magdy Younes.
Application Number | 20090107502 12/318555 |
Document ID | / |
Family ID | 30000723 |
Filed Date | 2009-04-30 |
United States Patent
Application |
20090107502 |
Kind Code |
A1 |
Younes; Magdy |
April 30, 2009 |
Method and device for monitoring and improving patient-ventilator
interaction
Abstract
Method and apparatus for non-invasively determining the time
onset (T.sub.onset) and end (T.sub.end) of patient inspiratory
efforts. A composite pressure signal is generated comprising the
sum of an airway pressure signal, a gas flow pressure signal
obtained by applying a gain factor (K.sub.f) to a signal
representing gas flow rate and a gas volume pressure signal
obtained by applying a gain factor (K.sub.v) to a signal
representing volume of gas flow. K.sub.f and K.sub.v values are
adjusted to result in a desired linear trajectory of composite
pressure signal baseline in the latter part of the exhalation
phase. The current composite pressure signal is compared with (i)
selected earlier composite pressure signal values and/or (ii) value
expected at current time based on extrapolation of composite
pressure signal trajectory at specified earlier times and/or (iii)
the current rate of change in the composite pressure signal with a
selected earlier rates of change. The differences obtained by the
comparison are compared with selected threshold values. T.sub.onset
is identified when at least one of the differences exceeds the
threshold values.
Inventors: |
Younes; Magdy; (Toronto,
CA) |
Correspondence
Address: |
SIM & MCBURNEY
330 UNIVERSITY AVENUE, 6TH FLOOR
TORONTO
ON
M5G 1R7
CA
|
Family ID: |
30000723 |
Appl. No.: |
12/318555 |
Filed: |
December 31, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10606751 |
Jun 27, 2003 |
7484508 |
|
|
12318555 |
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Current U.S.
Class: |
128/204.23 |
Current CPC
Class: |
A61M 16/026 20170801;
A61M 2016/0027 20130101; A61M 2230/46 20130101; A61M 2016/0021
20130101; A61M 2016/0039 20130101; A61M 2016/0042 20130101 |
Class at
Publication: |
128/204.23 |
International
Class: |
A61M 16/00 20060101
A61M016/00 |
Claims
1. (canceled)
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20. A method for cycling off the inflation phase of a mechanical
ventilator comprising: measuring the average interval between
patient inspiratory efforts in a patient in a suitable number of
elapsed breaths (T.sub.TOT) with said average being updated at
suitable intervals; identifying onset of current inspiratory
effort; monitoring time from said onset of inspiratory effort; and
generating a signal that causes the ventilator to cycle off when
time elapsed since onset of inspiratory effort exceeds a specified
fraction of T.sub.TOT.
21. The method of claim 20 wherein the time to generate a signal to
cycle off the ventilator is calculated from the trigger time of
current ventilator cycle plus a specified fraction of
T.sub.TOT.
22. A method for cycling off the inflation phase of a ventilator in
pressure support ventilation comprising: measuring the interval
between successive inspiratory efforts in a suitable number of
elapsed breaths (T.sub.TOT); measuring inspiratory flow rate at
specified times in those elapsed breaths which triggered ventilator
cycles, said specified times corresponding to a specified fraction
of the T.sub.TOT, measured from the onset of inspiratory effort of
said each breath or from the trigger time of the ventilator;
calculating the average of the flow values obtained at said
specified times in said elapsed breaths; and generating a signal
that causes the ventilator to cycle off when inspiratory flow in
the current inflation phase decreases below said average flow
value.
23. The method of claim 1 wherein results concerning patient
ventilator interaction are displayed, such results including
displays of at least one of the composite pressure signal itself,
T.sub.onset and T.sub.end markers, and trigger delay, cycling-off
errors, patient respiratory rate, number and frequency of
ineffective efforts, and frequency and duration of central apneas,
desirable duration of inflation phase, and flow at a specified
fraction of T.sub.TOT of the patient in the pressure support
ventilation mode.
24. (canceled)
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38. A device for estimating a desirable duration of inflation phase
of a ventilator, comprising: circuitry to identify inspiratory
efforts of a patient; means to calculate the time difference
between patient inspiratory efforts (patient T.sub.TOT); means for
displaying a value corresponding to a specified fraction of patient
T.sub.TOT, said specified fraction being a user input or a default
value between 0.3 and 0.5.
39. The device of claim 38 where a signal is generated to cycle off
the inflation phase of the ventilator when said desirable duration
has elapsed after ventilator triggering.
40. The device of claim 38 wherein a signal is generated to cycle
off the inflation phase of the ventilator when said desirable
duration has elapsed after onset of inspiratory effort in current
breaths or after a point intermediate between onset of effort and
ventilator triggering.
41. The device of claim 39 wherein a user input is provided for
inputting patient T.sub.TOT or its reciprocal, patient respiratory
rate, and said input is used by the device, in lieu of
device-determined patient T.sub.TOT, to determine desirable
duration of inflation phase.
42. A device for determining the desirable inspiratory flow
threshold for terminating inflation cycles in the pressure support
ventilation mode comprising: circuitry for estimating desirable
duration of inflation phase of the ventilator; means to measure
inspiratory flow in recently elapsed breaths after said desirable
duration has elapsed from the ventilator's trigger time, or from
the onset of inspiratory effort preceding triggered breaths, or
from a specified point in between the latter two points; and means
for displaying the value of said measured flow.
43. The device of claim 42 wherein the value of said measured flow
is communicated to cycling mechanism of the ventilator to effect
termination of the inflation phase when said measured flow, or a
reasonable approximation thereof, is reached during the inflation
phase.
44. The device of claim 24 wherein values relevant to patient
ventilator interaction are calculated and displayed, such values
including displays of at least one of the composite pressure signal
itself, T.sub.onset and T.sub.end markers and displays or outputs
indicating trigger delay, cycling-off errors, patient respiratory
rate, number and frequency of ineffective efforts, frequency and
duration of central apneas, desirable duration of inflation phase,
and flow at a specified fraction of patient's T.sub.TOT in the
pressure support ventilation mode.
45. The device of claim 24 wherein the output of the device is used
for closed-loop control of ventilator settings.
46. The device of claim 24 wherein functions executed by electrical
circuitry are executed in whole or in part by digital techniques.
Description
REFERENCE TO RELATED APPLICATION
[0001] This application claims priority under 35 USC 119(e) from
U.S. Provisional Patent Application No. 60/391,594 filed Jun. 27,
2002.
FIELD OF INVENTION
[0002] This invention relates to assisted mechanical
ventilation.
BACKGROUND TO THE INVENTION
[0003] With assisted ventilation (e.g. assist volume cycled
ventilation, pressure support ventilation and proportional assist
ventilation) ventilator cycles are triggered by the patient and are
intended to coincide with patient's inspiratory effort. In
practice, however, the ventilator cycle never begins at the onset
of patient's inspiratory effort (trigger delay) and the end of the
ventilator's inflation phase only rarely coincides with the end of
inspiratory effort (cycling-off errors). FIG. 1 provides an
example. The bottom channel is transdiaphragmatic pressure
(measured by esophageal and gastric catheters) and reflects true
patient inspiratory effort. As may be seen, ventilator cycle was
triggered several hundred milliseconds after onset of effort
(interval between vertical lines) and the inflation cycle continued
well beyond the effort. In fact, the ventilator was cycling almost
completely out-of-phase with the patient. Trigger delay is often so
marked that some efforts completely fail to trigger the ventilator
(ineffective efforts, e.g. third effort, FIG. 1). A more advanced
form of non-synchrony is shown in FIG. 2. In this case, the
inflation cycle of the ventilator extends over two patient cycles.
There are, accordingly, two inspiratory efforts within a single
inflation phase and there is an additional ineffective effort
during the ventilator's expiratory phase. The arrows in FIG. 2
indicate the location of the extra patient efforts that did not
trigger corresponding ventilator cycles.
[0004] Non-synchrony between patient and ventilator is extremely
common. Leung et al found that, on average, 28% of patient's
efforts are ineffective (Leung P, Jubran A, Tobin M J (1997).
Comparison of assisted ventilator modes on triggering, patient
effort, and dyspnea. Am J Respir Crit Care Med 155:1940-1948).
Considering that ineffective efforts are the extreme manifestation
of non-synchrony, less severe, yet substantial (e.g. first two
breaths, FIG. 1), delays must occur even more frequently.
Non-synchrony is believed to cause distress, leading to excessive
sedation and sleep disruption, as well as errors in clinical
assessment of patients since the respiratory rate of the ventilator
can be quite different from that of the patient. Monitoring
respiratory rate is a fundamental tool for monitoring critically
ill patients on ventilators.
[0005] In current mechanical ventilators, triggering occurs when
flow becomes inspiratory (i.e. >0) and exceeds a specified
amount, or when airway pressure decreases below the set PEEP
(positive end-expiratory pressure) level by a specified amount.
Trigger delay has two components. One component is related to
ventilator trigger response and sensitivity. Thus, if the response
of the ventilator is poor, triggering may not occur immediately
when the triggering criteria are reached. Alternatively, the
threshold for triggering may be set too high by the user. The
component of trigger delay attributable to ventilator response and
sensitivity is given by the interval between zero flow crossing
(arrow, FIG. 1) and triggering (second vertical line). The response
of modern ventilators has improved substantially over the past
several years such that it is difficult to effect further
improvements in this respect, and this invention does not
contemplate any such improvements. This component of trigger delay
can, however, still be excessive if the user sets an unnecessarily
high threshold. This setting may be because of lack of sufficient
expertise, or because there was excessive baseline noise at some
point, which necessitated a high threshold to avoid
auto-triggering. The threshold then remains high even after
disappearance of the noise.
[0006] The second component of trigger delay is the time required,
beyond the onset of inspiratory effort (T.sub.onset), for
expiratory flow to be reduced to zero (interval between first
vertical line and the arrow, FIG. 1). This delay is related to the
fact that expiratory resistance is usually high in ventilated
patients and expiratory time is frequently too short to allow lung
volume to return to FRC (functional residual capacity) before the
next effort begins. At T.sub.onset, therefore, elastic recoil
pressure is not zero (DH, dynamic hyperinflation). Inspiratory
effort must first increase enough to offset the elastic recoil
pressure associated with DH before flow can become inspiratory,
and/or before P.sub.aw (airway pressure) decreases below PEEP, in
order to trigger the ventilator. By identifying the true
T.sub.onset, which is one aspect of the current invention, this
component of trigger delay (usually the largest component, seen,
for example, FIG. 1) can be essentially eliminated.
[0007] Cycling-off errors result from the fact that, except with
Proportional Assist Ventilation, current ventilator modes do not
include any provision that links the end of ventilator cycle to end
of the inspiratory effort of the patient. In the most common form
of assisted ventilation, Volume Cycled Ventilation, the user sets
the duration of the inflation cycle without knowledge of the
duration of patient's inspiratory effort. Thus, any agreement
between the ends of ventilator and patient inspiratory phases is
coincidental. With the second most common form, Pressure Support
Ventilation, the inflation phase ends when inspiratory flow
decreases below a specified value. Although the time at which this
threshold is reached is, to some extent, related to patient effort,
it is to the largest extent related to the values of passive
resistance and elastance of the patient. In patients in whom the
product [resistance/elastance], otherwise known as respiratory time
constant, is high, the ventilator cycle may extend well beyond
patient effort, while in those with a low time constant the cycle
may end before the end of patient's effort (Younes M (1993)
Patient-ventilator interaction with pressure-assisted modalities of
ventilatory support. Seminars in Respiratory Medicine 14:299-322;
Yamada Y, Du H L (2000) Analysis of the mechanisms of expiratory
asynchrony in pressure support ventilation: a mathematical
approach. J Appl Physiol 88:2143-2150).
[0008] In U.S. Pat. No. 6,305,374 B1, an approach is described to
identify the onset and end of patient's inspiratory effort during
non-invasive bi-level positive pressure ventilation (BiPAP). This
approach relies exclusively on the pattern of flow waveform to make
these identifications. Thus, current values of flow are compared
with an estimated value based on projections from preceding flow
pattern. If the difference exceeds a preset amount, a phase switch
is declared. While this method may yield reasonably accurate
results in the intended application (treatment of obstructive sleep
apnea patients with non-invasive BiPAP), a number of considerations
suggest that its use in critically ill, intubated, ventilated
patients may not provide accurate results:
[0009] 1) Implicit to the use of flow as a marker of respiratory
muscle pressure output is the assumption that flow pattern reflects
changes in alveolar pressure inside patient's lung. This is where
respiratory muscle pressure is exerted. This assumption, however,
is true only if airway pressure is constant. Since airway pressure
is one of the two pressure values that determine flow (flow-(airway
pressure-alveolar pressure)/resistance), it is clear that changes
in airway pressure can alter flow even if there is no change in
respiratory muscle pressure. In non-invasive bi-level support,
airway pressure, one of the two pressure values that determine
flow, is reasonably constant during both inspiration and
expiration, even though the absolute level is different in the two
phases. If one of the two pressure values is constant during a
given phase, it is reasonable to assume that changes in flow during
that phase reflect changes in the other pressure, namely alveolar
pressure. This condition does not apply in intubated, mechanically
ventilated patients. In most modern intensive care ventilators,
airway pressure is actively controlled during expiration through
adjustments of the PEEP/exhalation valve mechanism. The pattern of
such active changes in airway pressure during expiration varies
from one ventilator brand to another and in the same ventilator
from time to time depending on the state of the PEEP/exhalation
valve mechanism. Under these conditions, changes in flow trajectory
during expiration cannot be assumed to reflect changes in alveolar
pressure trajectory. Likewise, during inspiration airway pressure
is far from being constant, regardless of the mode used. Thus,
changes in inspiratory flow profile cannot be used to reflect
similar changes in alveolar pressure. The use of flow to infer end
of effort during the inflation phase is accordingly not
plausible.
[0010] 2) When passive elastance (E) and resistance (R) are
constant over the entire tidal volume range, the product R/E, or
respiratory time constant, is also constant over the entire period
of expiration. Because the time constant governs the pattern of
lung emptying, a constant R/E produces a predictable exponential
flow pattern in the passive system. With a predictable pattern it
is possible to make forward extrapolations, or predictions, for the
sake of identifying a deviation from the expected passive
behaviour. Such deviation may then be used, with reasonable
confidence, to infer the development of an additional active force,
such as the onset of inspiratory muscle effort. When E and R are
not constant throughout the breath, R/E may change from time to
time causing changes in flow trajectory (.DELTA.flow/.DELTA.t) that
are not related to muscle pressure. Under these conditions,
deviation in .DELTA.flow/.DELTA.t from previous values cannot
reliably signify a change in pressure generated by respiratory
muscles. Patients with obstructive sleep apnea, the intended
population of U.S. Pat. No. 6,305,374 B1, have generally normal
lungs; R and E are expected to be constant over the tidal volume
range, particularly when expiratory airway pressure is higher than
atmospheric (i.e. the usual case when BiPAP is applied). In
critically ill, intubated ventilated patients, this is not the
case. Resistance is not constant, primarily because these patients
are intubated and the resistance of the endotracheal tube is
flow-dependent (the higher the flow, the higher the resistance).
The relation between resistance and flow varies from one tube to
the other. Furthermore, tidal volume in these patients often
extends into the volume range where elastance is not constant.
Thus, as the lung is emptying, either or both elastance and
resistance may be changing, causing changes in respiratory time
constant during the same expiration. Under these conditions,
changes in flow trajectory need not reflect changes in respiratory
muscle pressure. This considerably decreases the sensitivity and
specificity of flow pattern as a marker of inspiratory effort.
[0011] 3) Changes in respiratory muscle pressure (P.sub.mus) are
not exclusively used to change flow. According to the equation of
motion, specifically applied to intubated patients:
P.sub.mus=Volume*E+Flow*K.sub.1+(Flow*absolute
flow*K.sub.2)-P.sub.aw Equation 1
[0012] Where, E is passive respiratory system elastance, K.sub.1 is
the laminar component of passive respiratory system resistance,
K.sub.2 is the resistance component related to turbulence (mostly
in the endotracheal tube), and P.sub.aw is airway pressure which is
determined by the pressure at the exhalation/PEEP valve
(P.sub.valve), flow and R.sub.ex, that is resistance of the
exhalation tubing (P.sub.aw=P.sub.valve-flow*R.sub.ex). In this
equation expiratory flow is negative. When P.sub.mus changes, as at
T.sub.onset, the flow trajectory should change. However, a change
in flow trajectory also results in changes in volume and P.sub.aw
trajectories. According to Equation 1, these changes will oppose
the change in flow. For example, if expiratory flow decreases at a
faster rate, volume decreases at a slower rate than in the absence
of P.sub.mus. At any instant after T.sub.onset, elastic recoil
pressure, which is related to volume, is higher, and this promotes
a greater expiratory flow. The same can be said for the effect of
changes in flow trajectory on P.sub.aw trajectory; a lower
expiratory flow decreases P.sub.aw, which promotes more expiratory
flow. How much of the change in P.sub.mus is used to change the
flow trajectory depends on the magnitude of the opposing forces. In
particular, a higher passive elastance and/or a higher R.sub.ex
tends to reduce the fraction of the change in P.sub.mus used to
change flow trajectory. Furthermore, for a given P.sub.mus expended
to change the flow trajectory, the actual change in trajectory is
determined by resistance (i.e. K.sub.1 and K.sub.2). When E,
R.sub.ex K.sub.1 and K.sub.2 are all low, a modest change in
dP.sub.mus/dt results in a sharp change in flow trajectory. As
these characteristics become more abnormal, the change in flow
trajectory, for a given dP.sub.mus/dt, progressively is attenuated.
FIG. 3 illustrates this in a computer simulation.
[0013] In the example of FIG. 3, respiratory muscles were inactive
in the first second of expiration (as they usually are). This is
represented by P.sub.mus of zero (lower panel). At 1.0 sec an
inspiratory effort begins. P.sub.mus rises at a rate of 10
cmH.sub.2O/sec, representative of a normal respiratory drive. The
three flow waveforms represent, from below upwards, progressively
increasing values of K.sub.1, K.sub.2, E and R.sub.ex. The values
used in the lowest waveform are those of a patient with normal
passive elastance and resistance, intubated with a large
endotracheal tube (#9 tube, K.sub.2=3), and exhalation tubing with
a low resistance (R.sub.ex=2). The onset of effort results in a
sharp change in the flow trajectory that can be readily detected
within a very short time after T.sub.onset.
[0014] The middle waveform (FIG. 3) was generated with values
representing the average intensive care patient on mechanical
ventilation. Both passive K.sub.1 and passive E are higher than
normal, K.sub.2 is that of a #8 endotracheal tube, the most common
size used, and the exhalation tubing has a moderate (average)
resistance. Note that the change in flow trajectory is considerably
less pronounced. An experienced eye, with the benefit of hindsight
(i.e. observing the flow waveform for a substantial period after
P.sub.mus started), may be able to tell that a change in trajectory
occurred at 1.0 sec. However, it is not possible to prospectively
identify that a trajectory change took place in a timely manner,
for the sake of triggering the ventilator. Prospective
identification of a trajectory change requires comparison between
current and previous .DELTA.flow/.DELTA.t values, or between
current flow values and values expected based on forward
extrapolation of the preceding flow pattern (e.g. dashed lines,
FIG. 3). There is always uncertainty with extrapolation,
particularly with non-linear functions where the exact function is
not known and, even more so, when the signal is noisy, as the flow
signal commonly is (due to cardiac artefacts or secretions).
Comparison of current and previous .DELTA.flow/.DELTA.t is also
fraught with uncertainties when the rate may change for reasons
other than respiratory muscle action (see #1 and #2, above). Thus,
a wide difference (trigger threshold) must be specified, between
current and projected flow, or between current and previous
.DELTA.flow/.DELTA.t, before a trajectory change can be identified
with confidence. Otherwise, false triggering will occur frequently.
When the change in flow trajectory is small, a longer interval must
elapse before the threshold separation is achieved. It can be seen
from the middle flow waveform that a conservative flow separation
(between actual and projected flow) of 0.2 l/sec would not be
reached until after flow became inspiratory. Thus, in the average
mechanically ventilated patient the use of flow trajectory to
identify T.sub.onset is not likely to result in a significant
improvement over the current approach of waiting for flow to become
inspiratory.
[0015] With more severe mechanical abnormalities (top waveform,
FIG. 3), the change in flow trajectory is even more subtle. Even an
experienced eye, with the benefit of hindsight, cannot distinguish
between a true trajectory change and some flow artefact. Clearly,
with a much stronger effort a flow trajectory change may be
identifiable before flow becomes inspiratory. However, when
patients have vigorous inspiratory efforts, there is no significant
trigger delay even with current triggering techniques.
[0016] In summary, the use of flow to identify respiratory phase
transitions is entirely unsuitable for identification of
inspiratory to expiratory transitions during mechanical ventilation
in critically ill patients (because of the highly variable P.sub.aw
during inflation), and has poor sensitivity and specificity for
identifying expiratory to inspiratory transitions in these patients
because of the frequent use of active exhalation valves, the
presence of variable time constant during expiration and the often
marked abnormalities in elastance and resistance.
SUMMARY OF INVENTION
[0017] In one aspect, the present invention provides a method for
detecting the onset of inspiratory effort (T.sub.onset) in a
patient on mechanical ventilation, comprising the steps of:
[0018] (a) monitoring airway pressure, rate of gas flow, and volume
of gas flow of the patient;
[0019] (b) applying a gain factor (K.sub.f) to the signal
representing rate of gas flow to convert the gas flow signal into a
gas flow pressure signal;
[0020] (c) applying a gain factor (K.sub.v) to the signal
representing volume of gas flow to convert the gas volume signal
into a gas volume pressure signal;
[0021] (d) generating a composite pressure signal (signal)
comprising the sum of airway pressure signal, gas flow pressure
signal, and gas volume pressure signal, with all signals, having a
suitably adjusted polarity;
[0022] (e) adjusting K.sub.f and K.sub.v to result in a desired
linear trajectory of composite pressure signal baseline in the
latter part of the exhalation phase; [0023] (f) comparing (i) the
current composite pressure signal values with selected earlier
composite pressure signal values; and/or [0024] (ii) the current
composite pressure signal values with values expected at current
time based on extrapolation of composite pressure signal trajectory
at specified earlier times, and/or [0025] (iii) the current rate of
change in the composite pressure signal with a selected earlier
rate of change in the composite pressure signal;
[0026] (g) comparing differences obtained from such comparison(s)
made in step (f) with selected threshold values; and
[0027] (h) identifying T.sub.onset when at least one of the
differences exceeds the threshold values.
[0028] The composite pressure signal may contain a fourth
component, consisting of the square of the rate of gas flow to
which a gain factor (K.sub.f2) is applied to convert the fourth
signal to a pressure signal. K.sub.f2 may also be used to adjust
the trajectory of the composite pressure signal baseline in the
latter part of the exhalation phase. K.sub.f2 may be assigned a
value corresponding to the K.sub.2 constant of an endotracheal tube
in the patient. The values of K.sub.v, K.sub.f and/or K.sub.f2 may
be adjusted to result in a specified slope or pattern of the
composite pressure signal during part or all of the expiratory
phase.
[0029] A default value of K.sub.f may be used while the value of
K.sub.v is adjusted to obtain a desired baseline composite pressure
signal trajectory. Alternatively, a default value of K.sub.v is
used while the value of K.sub.f is adjusted to obtain a desired
baseline composite pressure signal trajectory.
[0030] The K.sub.f or K.sub.v value used may be a known or
estimated value of the respiratory system resistance or elastance,
respectively, of the patient.
[0031] The current composite pressure signal value may be compared
with the composite pressure signal value at the most recent point
where the composite pressure signal began a new rising phase and
T.sub.onset is identified when the calculated difference exceeds a
set threshold value.
[0032] T.sub.onset detection may be precluded in the early part of
the exhalation phase.
[0033] The amplitude of the composite pressure signal may be
monitored through the inspiratory phase and the end of inspiratory
effort (T.sub.end) is identified from a reduction in signal
amplitude or signal slope below a specified value, which may be a
specified fraction of the highest value obtaining during the
inspiratory phase. T.sub.end detection may be precluded in the
early part of the inflation phase. The generated signals
corresponding to T.sub.onset may be used to trigger ventilation
cycles and/or signals corresponding to T.sub.end may be used to
cycle off ventilation cycles.
[0034] In another aspect of the invention, there is provided a
method for detecting the onset of inspiratory effort (T.sub.onset)
in a patient on mechanical ventilation, comprising the steps
of:
[0035] (a) monitoring airway pressure and rate of gas flow of the
patient,
[0036] (b) applying a gain factor (K.sub.f) to the signal
representing rate of gas flow to covert the gas flow signal into a
gas flow pressure signal,
[0037] (c) generating a composite pressure signal comprising the
sum of airway pressure signal and the gas flow pressure signal,
[0038] (d) comparing (i) the current composite pressure signal
values with values expected based on extrapolation of composite
pressure signal trajectory at specified earlier times, and/or
[0039] (ii) the current rate of change of composite pressure signal
with a selected earlier rate of change of composite pressure
signal,
[0040] (e) comparing differences obtained from such comparison(s)
made in step (d) with selected threshold values, and
[0041] (f) identifying T.sub.onset when at least one of the
differences exceeds said threshold values.
[0042] In this aspect of the invention, the composite pressure
signal may incorporate a third component consisting of the square
of the rate of gas flow, to which a gain factor (K.sub.f2) is
applied to convert the third signal to a pressure signal. The
selected K.sub.f may be a known or assumed value of respiratory
system resistance.
[0043] The generated signal representing T.sub.onset may be used to
trigger ventilation cycles.
[0044] The present invention further includes, methods for
determining a suitable threshold value for identifying the onset of
inspiratory effort from the composite pressure signal obtained
according to the procedures described above.
[0045] In one such method, suitable for use where the composite
pressure signal includes the sum of the airway pressure signal, gas
flow pressure signal and gas volume pressure signal, and,
optionally, the fourth component, comprises:
[0046] monitoring the composite pressure signal over suitable
intervals preceding onset of inspiratory effort, in a suitable
number of elapsed breaths;
[0047] identifying peaks and troughs in the composite pressure
signal over the duration of the intervals;
[0048] measuring the changes in signal amplitude between successive
peaks and troughs, the amplitudes reflecting the range of
amplitudes of noise included in the composite pressure signal;
and
[0049] determining from the detected range of noise amplitude, a
value that exceeds the prevailing noise value, such value then
being used prospectively to distinguish between true inspiratory
efforts and noise.
[0050] Another such method, suitable for use where the composite
pressure signal includes the sum of the airway pressure signal, gas
flow pressure signal and gas volume pressure signal, and
optionally, the fourth component, or where the composite pressure
signal includes the sum of the airway pressure signal and the gas
flow pressure signal, and optionally, the third component,
comprising:
[0051] monitoring the composite pressure signal over suitable
interval preceding onset of inspiratory effort in a suitable number
of elapsed breaths;
[0052] determining slope of the composite pressure signal in
successive subintervals within the intervals;
[0053] measuring the range of slope in the subintervals, such range
reflecting the range of slope change in composite pressure signal
related to noise; and
[0054] determining from the detected range of slope changes, a
difference in slope that exceeds the prevailing noise level, the
resulting value then being used prospectively to distinguish
between changes in composite pressure signal slope due to
inspiratory efforts and those due to composite pressure signal
noise.
[0055] An alternative to the latter method comprises:
[0056] monitoring the composite pressure signal over suitable
intervals preceding the onset of inspiratory effort, in a suitable
number of elapsed breaths;
[0057] comparing signal amplitude at discrete points within such
intervals with values predicted to occur at such times from the
signal pattern in previous intervals, the difference in signal
amplitude reflecting the range of difference related to composite
pressure signal noise; and
[0058] determining from the detected range of differences, a value
that exceeds the prevailing noise level, such value then being used
prospectively to identify differences between current and predicted
values that reflect true inspiratory effort.
[0059] In another aspect of the present invention, there is
provided a method for cycling off the inflation phase of a
mechanical ventilator, which comprises:
[0060] measuring the average interval between successive
inspiratory efforts in a patient in a suitable number of elapsed
breaths (T.sub.TOT);
[0061] identifying onset of inspiratory effort by utilizing any of
the procedures provided in accordance with the present invention or
otherwise;
[0062] monitoring the time from the onset of inspiratory effort;
and
[0063] generating a signal that causes the ventilator to cycle off
when time elapsed since onset of inspiratory effort exceeds a
specified fraction of T.sub.TOT.
[0064] The time to generate a signal to cycle off the ventilator
may be calculated from the trigger time of current ventilation
cycle plus a specified fraction of T.sub.TOT.
[0065] In a further aspect of the present invention, there is
provided a method for cycling off the inflation phase of a
ventilator in pressure support ventilation, comprising:
[0066] measuring the interval between successive inspiratory
efforts in a suitable number of elapsed breaths (T.sub.TOT);
[0067] measuring inspiratory flow rate at specified times in the
elapsed breaths which triggered ventilator cycles, the specified
times corresponding to a fraction of the T.sub.TOT, measured from
the onset of inspiratory effort of each breath or from the trigger
time of the ventilator;
[0068] calculating the average of the flow values obtained at such
specified times in the elapsed breaths; and
[0069] generating a signal that causes the ventilator to cycle off
when inspiratory flow in the current inflation phase decreases
below said average flow value.
[0070] The results concerning patient ventilator interaction may be
displayed in suitable format, including but not limited to a
monitor, digital or electrical output ports, or printed material.
Such results may include, but not limited to, display of the
composite pressure signal, T.sub.onset and T.sub.end markers and
displays regarding trigger delay, cycling-off errors, patient
respiratory rate, number and frequency of ineffective efforts, and
frequency and duration of central apneas, desirable duration of
inflation phase, and flow at a specified fraction of T.sub.TOT of
the patient in the pressure support ventilation mode.
[0071] In accordance with another aspect of the present invention,
there is provided an apparatus for detecting the onset of
inspiratory effort (T.sub.onset) in a patient on mechanical
ventilation, comprising:
[0072] circuitry for measuring airway pressure, rate of gas flow
and volume of gas flow of the patient;
[0073] amplifier to apply a gain factor (K.sub.f) to the signal
representing rate of gas flow to convert the signal into a gas flow
pressure signal;
[0074] amplifier to apply a gain factor (K.sub.v) to the signal
representing volume of gas flow to convert the signal into a gas
volume pressure signal;
[0075] summing amplifier that generates a composite pressure signal
comprising the sum of airway pressure signal, the gas flow pressure
signal and the gas volume pressure signal, with all signals having
suitably adjusted polarity;
[0076] means to permit adjustment of K.sub.f and K.sub.v to provide
a desired trajectory of composite pressure signal baseline in the
latter part of the exhalation phase;
[0077] circuitry to direct the composite pressure signal to a
T.sub.onset identification circuitry during a suitable period in
the expiratory phase, the identification circuitry comprising
circuitry to detect a change in trajectory; and
[0078] means for generating a signal corresponding to T.sub.onset
when measured change in composite pressure signal trajectory
exceeds a specified threshold.
[0079] In the device of the invention, an additional signal may be
generated to be summed by the summing amplifier being generated by
multiplying the flow signal by the absolute value of the flow
signal and applying a gain factor (K.sub.f2) to the resulting
square flow signal using an amplifier and K.sub.f2 is also used to
adjust the trajectory of the composite pressure signal baseline in
the latter part of the exhalation phase. K.sub.f2 may be assigned a
value corresponding to the K.sub.2 constant of the endotracheal
tube in place in the patient.
[0080] The K.sub.f value may be fixed at a default value while
adjustment of signal trajectory is made using K.sub.v and/or
K.sub.f2. Alternatively, K.sub.v is fixed at a default value while
adjustment of signal trajectory is made using K.sub.f and/or
K.sub.f2.
[0081] In one embodiment of the invention, the summing amplifier
input related to volume of flow is omitted.
[0082] The device provided herein may include circuitry that
precludes T.sub.onset identification during an adjustable period
after the end of the inflation phase of the ventilator.
[0083] The T.sub.onset identification circuitry may comprise
circuitry to obtain the rate of change of composite pressure signal
amplitude and to obtain the difference between the current rate of
change and the rate of change of the composite pressure signal
amplitude at a specified earlier time and to generate a T.sub.onset
signal when the difference exceeds a set threshold value.
[0084] The T.sub.onset identification circuitry may comprise
circuitry to measure the difference between the current composite
pressure signal amplitude and the composite pressure signal
amplitude at a specified earlier time and to generate a T.sub.onset
signal when the difference exceeds a set threshold value.
[0085] In the device of the invention, K.sub.v and/or K.sub.f
and/or K.sub.f2 may be adjusted to produce a horizontal or slightly
downward sloping composite pressure signal baseline in the latter
part of expiration and the T.sub.onset identification circuitry may
comprise circuitry to measure the difference between current
composite pressure signal amplitude and composite pressure signal
amplitude at the most recent point where the composite pressure
signal began rising and to generate a T.sub.onset signal when the
difference exceeds a set threshold value.
[0086] The composite pressure signal may be gated to circuitry to
identify end of inspiratory effort (T.sub.end), such circuitry
comprising:
[0087] circuitry to identify the highest amplitude (peak) of the
composite pressure signal reached during the current inspiratory
effort;
[0088] circuitry to detect when amplitude of the composite pressure
signal decreases below a specified value beyond the time at which
the peak occurred; and
[0089] circuitry to generate a signal corresponding to T.sub.end
when the amplitude of the composite pressure signal decreases below
the specified value, which may be a specified fraction of the peak
amplitude of the composite pressure signal. Circuitry may be
provided to preclude detection of T.sub.end during a specified
period following ventilator triggering.
[0090] Signal corresponding to T.sub.onset maybe used to trigger
ventilator cycles and/or signal corresponding to T.sub.end may be
used to cycle off inflation phases of the composite pressure
signal.
[0091] The output of the device may be used for closed-loop control
of ventilation setting. Functions executed by electrical circuitry
may be executed in whole or in part by digital techniques.
[0092] In a further aspect of the present invention, there is
provided a device for estimating a desirable duration of the
inflation phase of a ventilator, comprising:
[0093] circuitry to identify inspiratory efforts of the patient,
which may be a device according to the invention or by other
suitable circuitry;
[0094] means to calculate the time difference between patient
inspiratory efforts (patient T.sub.TOT); and
[0095] means for displaying a value corresponding to a specified
fraction of patient T.sub.TOT, such specified fraction being a user
input or a default value between 0.3 and 0.5.
[0096] In this device, a signal may be generated to cycle off the
inflation phase of the ventilator when the desirable duration has
lapsed after ventilator triggering.
[0097] A signal may be generated to cycle off the inflation phase
of the ventilator when the desirable duration has elapsed after
onset of inspiratory effort in current breaths or after a point
intermediate between onset of effort and ventilator triggering.
[0098] A user input may be provided for inputting patient T.sub.TOT
or its reciprocal, patient respiratory rate, and the input then is
used by the device, in lieu of device-determined patient T.sub.TOT,
to determine desirable duration of inflation phase.
[0099] In an additional aspect of the invention, there is provided
a device for determining the desirable inspiratory flow threshold
for terminating inflation cycles in the pressure support
ventilation mode, comprising:
[0100] circuitry for estimating desirable duration of inflation
phase of the ventilator, by using the device provided herein or by
any other suitable alternative;
[0101] means for measuring inspiratory flow in recently elapsed
breaths after the desirable duration has elapsed from the
ventilator trigger time, or from the onset of inspiratory effort
preceding triggered breaths, or from a specified point in between
the two points; and
[0102] means for displaying the value of said measured flow.
[0103] In such device, the value of the measured flow may be
communicated to the cycling mechanism of the ventilator to effect
termination of the inflation phase when the measured flow, or a
reasonable approximate thereof, is reached during the inflation
phase.
[0104] The values relating to patient ventilator interaction
determined in the devices provided herein may be calculated and
displayed in suitable format, including but not limited to a
monitor, digital or electrical output ports. The values may be any
of those discussed above.
[0105] The present invention, therefore, concerns a novel method
and apparatus to, non-invasively, determine the true onset
(T.sub.onset) and end (T.sub.end) of patient's inspiratory efforts.
Such method/device can be used simply as a monitor, informing the
user of the presence and magnitude of trigger delays, ineffective
efforts and cycling-off errors. The user can then take appropriate
action to reduce the non-synchrony. Alternatively, the
method/device can be coupled with the cycling mechanisms of the
ventilator, whereby onset and end of ventilator cycles are
automatically linked to onset and end of patient's efforts, thereby
insuring synchrony without intervention by the user.
[0106] One aspect of the current invention is to minimize the
cycling-off errors either by directly identifying the end of
patient's inspiratory effort or by insuring that the ventilator's
inflation phase does not extend beyond the physiologic limit of the
duration of inspiratory effort.
BRIEF DESCRIPTION OF DRAWINGS
[0107] FIG. 1 contains traces of airway pressure, flow and
diaphragm pressure for a patient on mechanical ventilation;
[0108] FIG. 2 contains further traces of airway pressure, flow and
diaphragm pressure for ventilator cycles;
[0109] FIG. 3 is a graphical representation of the effect of
variation in certain parameters on change in trajectory of flow
upon start of inspiration;
[0110] FIG. 4 is a graphical representation of the effect of
variation in certain parameters on change in trajectory of
composite pressure signal Z upon start of inspiration;
[0111] FIG. 5 contains traces of airway pressure, flow and
composite pressure signal Z calculated in accordance with the
invention;
[0112] FIG. 6 contains traces of airway pressure, flow, composite
pressure signal Z and diaphragm electrical activity, with the
signal Z tracing being generated from pressure, flow and volume
tracings;
[0113] FIG. 7 is a schematic representation of the generation of
pressure and flow signals;
[0114] FIG. 8 is a block diagram of one embodiment of a device
operating in accordance with the method of the invention;
[0115] FIG. 9 is a schematic representation of the digital
implementation of output functions;
[0116] FIG. 10 contains traces of composite pressure signal Z and
T.sub.onset integrator output;
[0117] FIGS. 11 and 12 show the electrical circuitry used in
apparatus of FIG. 8;
[0118] FIGS. 13 to 17 contain flow charts for the different
functions performed by the output microprocessor shown in FIG.
9;
[0119] FIG. 18 is a block diagram of one embodiment of a fully
digital device for carrying out the method of the invention;
and
[0120] FIGS. 19 to 21 contain flow charts for the different
functions performed by the fully digital device of FIG. 18.
DETAILED DESCRIPTION OF THE INVENTION
[0121] The present invention contemplates novel methods and devices
for specific and timely identification of respiratory phase
transitions within the patient for use in monitoring
patient-ventilator interaction or to effect switching of ventilator
cycles. These methods/devices represent a progression in complexity
that address the problems inherent in the prior art ventilation
procedures described above.
[0122] In the simplest of these methods, a signal is generated
(signal X) that incorporates changes in both the flow and airway
pressure (P.sub.aw) signals. Thus,
Signal X=(Flow*K.sub.f)-P.sub.aw Equation 2,
[0123] where, K.sub.f is a constant that converts flow to pressure.
K.sub.f may be an estimated or assumed value of patient's
resistance (including endotracheal tube). There are two advantages
to this approach: First, the signal becomes relatively immune to
changes in flow trajectory produced via changes in pressure at the
exhalation/PEEP valve mechanism (#1 in Background above). Thus, if
pressure at the exhalation/PEEP valve increased near the end of
expiration (to maintain PEEP), flow will decrease at a faster rate.
Without the P.sub.aw component, this effect may appear as an
inspiratory effort. With inclusion of P.sub.aw in the signal,
changes in flow and P.sub.aw tend to cancel out. The extent to
which this compensation is complete depends on how close K.sub.f is
to actual patient resistance. In the absence of a known value, a
default value may be used, for example 15 cmH.sub.2O/l/sec,
representing average resistance (including ET tube) in critically
ill, mechanically ventilated patients. With such a default value,
correction is not perfect, but the signal is more specific (than
flow) in reflecting T.sub.onset. Second, by including P.sub.aw in
the signal, the signal incorporates that component of P.sub.mus
that was dissipated against R.sub.ex (see #3 in Background). For
example, if P.sub.aw decreases at T.sub.onset (because of the lower
expiratory flow), this decrease is summed with the component
related to flow, resulting in a sharper change in signal
trajectory. With this approach, however, signal baseline prior to
inspiratory effort is not flat, but, as in the case of flow, rises
in a non-linear fashion. Forward extrapolation continues to be
required to identify phase transition. Thus, the uncertainty
associated with forward extrapolation is not eliminated but the
change in signal trajectory is sharper, resulting in a more timely
detection of T.sub.onset for the same selected detection threshold
(i.e. difference between actual and predicted signal required for
identification). Furthermore, this approach continues to be
unsuitable for detection of inspiration to expiration transitions
(T.sub.end).
[0124] A further improvement is achieved by incorporating a
component related to volume in the signal (signal Y). Thus:
Signal Y=Volume*K.sub.v+Flow*K.sub.f-P.sub.aw Equation 3,
[0125] where, K.sub.v is a factor that converts volume to pressure.
With this treatment, the increase in the flow term during
expiration (note that flow is negative) is offset by the decrease
in the volume term. This tends to linearize, and decrease the slope
of (flatten) the signal in the interval prior to T.sub.onset,
reducing the uncertainty associated with extrapolation, while the
change in trajectory at T.sub.onset is rendered more acute on
account of incorporating representation of all actions resulting
from the change in P.sub.mus (see #3 in Background). In the best
case scenario, where K.sub.v is identical to passive elastance,
K.sub.f is identical to passive resistance, and there are no
non-linearities in the passive pressure-flow and pressure-volume
relations, signal Y would be identical to the actual P.sub.mus
waveform, with a flat baseline and a crisp rising phase at
T.sub.onset (i.e. as in the P.sub.mus panel of FIG. 3). Under these
conditions, extrapolation is unnecessary, and phase transition is
identified when signal Y exceeds a set threshold above the baseline
value, to account for random baseline noise. Unfortunately,
however, precise determination of actual passive properties during
assisted ventilation is impossible, and there are non-linearities
in the pressure-flow and pressure-volume relations. These result in
some instability in baseline, necessitating the use of
extrapolation. It may be expected, however, that the transition
from baseline to active inspiration will be crisper after including
a volume component (see below).
[0126] A further improvement is achieved by allowing for
non-linearity in the pressure-flow relation. In mechanically
ventilated patients, the non-linear element is almost exclusively
due to endotracheal tube characteristics. Thus, a suitable
alternate approach is to partition the flow component in two parts,
one related to the endotracheal tube and the other related to a
laminar component of resistance (K.sub.f). Such signal is referred
to as signal Z. Thus:
Signal Z=Volume*K.sub.v+Flow*K.sub.f+(Flow*absolute
flow*K.sub.f2)-P.sub.aw Equation 4,
[0127] where K.sub.f2 may be the commercially available K.sub.2
value of the endotracheal tube in place. This treatment essentially
eliminates any artifactual baseline instability related to
non-linear pressure-flow behaviour, further reducing the need for
extrapolation and enhancing the crispness of the transition.
[0128] As indicated earlier, precise estimates of E and K.sub.1 are
impossible to obtain during assisted ventilation. Passive E and R
(including K.sub.1) may be available from earlier determinations in
which the patient was made passive. These values may be different
from the current values, either because the ventilation conditions
under which measurements were made were different, or true E and R
(i.e. K.sub.1) may have changed in the interim. Some techniques can
be used to estimate E and R during conventional assisted
ventilation, but these are not very reliable. An important issue,
therefore, is the impact of differences between the K.sub.v and
real E, and between K.sub.f and real resistance, on the baseline of
the generated signals and on the sharpness of the transition.
[0129] In FIG. 4, the same P.sub.mus waveform shown at the bottom
of FIG. 3 was used to generate flow, volume and P.sub.aw waveforms
using values representative of the average patient (K.sub.1=10,
K.sub.2=5.5, E=25, R.sub.ex=5, similar to the values used to
generate the middle flow panel of FIG. 3). Signal Z was then
generated from the resulting flow, volume and P.sub.aw waveforms
using inaccurate values of K.sub.v and K.sub.f (i.e. K.sub.v
different from real E and K.sub.f different from true K.sub.1).
Simulations were made with errors in either direction (over- or
underestimation) of a magnitude that reflects reasonable outside
limits of such errors in practice (i.e. E and K.sub.1 overestimated
by 100% or underestimated by 70%).
[0130] As may be expected, when there are no errors (i.e. K.sub.v=E
and K.sub.f=K1, middle line, FIG. 4), signal Z is identical to the
actual P.sub.mus waveform. However) when there are differences
between assumed values and actual values, the baseline; prior to
T.sub.onset, is neither flat nor linear. When K.sub.v is >E, or
K.sub.f is <K.sub.1 (upper two lines), baseline is sloping down.
Under these conditions, there is a qualitative change in direction
of signal Z at T.sub.onset of effort. Such a directional change can
be easily detected (e.g. by differentiating signal Z and looking
for the point at which the differentiated signal becomes positive).
However, when K.sub.v is <E, or K.sub.f is >K.sub.1 (bottom
two lines, FIG. 4), baseline is sloping up and T.sub.onset is
evident as a change in slope; a quantitative, as opposed to the
qualitative, difference observed with the opposite errors. To
identify inspiratory effort under these conditions, as in the case
of flow (FIG. 3), requires forward projection or extrapolation with
the attendant increase in uncertainty and the necessity to increase
trigger threshold. It should be noted, however, that with this
approach (i.e. using signal Z (or Y) as opposed to flow) the change
in trajectory is much sharper than in the case of flow (middle
line, FIG. 3), making it possible to identify inspiratory effort
sooner. It should also be noted that the upward slope of the
signal, once effort begins, is related to the K.sub.f value, being
higher when K.sub.f is higher than K.sub.1, and vice versa.
[0131] It follows that the use of known values of E and K.sub.1,
obtained from previous direct measurement, offers advantages over
the use of flow. However, under some conditions (i.e. baseline
sloping upward) extrapolation techniques (or comparisons between
current and previous rates of signal change) are required, and this
may delay detection of phase transition.
[0132] A further novel aspect of this invention is to completely
ignore patient values of E and K.sub.1 and to simply select empiric
values of K.sub.v and K.sub.f that result in a flat or slightly
downward sloping baseline in the latter part of expiration. It is
clear from FIG. 4 that, with respect to baseline pattern (i.e.
pattern prior to inspiratory effort), errors can be made to cancel
out. Thus, overestimation of E and overestimation of K.sub.1
produce opposite errors. If empiric values of K.sub.v and K.sub.f,
that may have no bearing on actual values, are used, the baseline
may be sloping up or down depending on the nature and magnitude of
errors. Even though one cannot tell which value is in error, or by
how much, it is always possible to obtain a flat baseline by
adjusting either K.sub.f or K.sub.v. For example, if using the
empiric values results in an upward sloping baseline, the baseline
can be made flat by increasing the empiric K.sub.v or decreasing
the empiric K.sub.f. If such adjustments result in a flat baseline
but some systematic non-linearities persist, these can be offset by
adjustments of the non-linear K.sub.f2 term, if signal Z is used,
resulting in a flat, and linear baseline. Under such conditions,
identification of T.sub.onset presents little difficulty. A
particularly suitable approach for generating signal Z is to use a
default K.sub.f value of 10 cmH.sub.2O/l/sec (15 if signal Y is
used) and adjust K.sub.v to obtain a flat signal baseline.
Alternatively, a default K.sub.v value (e.g. 25 cmH.sub.2O/l,
representing average elastance in ICU patients) is used and K.sub.f
is adjusted to obtain a flat signal baseline. The former approach
was found preferable by the inventor as it guarantees a fairly
brisk rate of signal rise at T.sub.onset. Adjustments of K.sub.v at
a set K.sub.f, or vice versa, can be implemented by the user
employing external inputs for K.sub.v and/or K.sub.f, with feedback
from a graphic display of the generated signal (signal Y or Z).
Alternatively, selection of the optimum K.sub.v and K.sub.f values
may be done automatically using appropriate software.
[0133] The above approach does not address the possibility of
non-linear passive pressure-volume relation in the tidal volume
range (i.e. non-constant elastance). When this is present, and it
is common in mechanically ventilated patients, the respiratory
system is stiffer in the higher part of the tidal range. When
K.sub.v, which is a constant, is adjusted to produce a flat or
slightly decreasing signal in the latter part of expiration the
signal is not flat in the early part of expiration. In the presence
of non-constant elastance (higher elastance at higher volumes) the
signal shows a rising phase in the early part of expiration that
continues until volume reaches the range of constant elastance.
This artifactual rising phase may cause false identification of a
new inspiratory effort. This problem is averted by "blinding" the
T.sub.onset detection circuitry to the signal during the early part
of expiration. This can be done, for example, by gating the signal
to the T.sub.onset detection circuitry only after a certain delay
from onset of expiratory flow (T.sub.onset window delay).
Alternatively, the T.sub.onset detection circuitry may continue to
detect T.sub.onset during this period but the resulting
identification is gated out during this period. Detection of these
false triggers can be easily recognized visually by their
consistent relation to end of ventilator cycle. The magnitude of
the delay (blinding or blanking period) can then be adjusted
accordingly. Alternatively, software algorithms can be developed to
detect triggering signals with a consistent relation to end of
ventilator cycle and automatically adjusting the width of the
window.
[0134] The approach of blinding the T.sub.onset detection circuitry
to the signal over a time zone close to ventilator cycling-off,
where flow is changing rapidly, also helps weed out false triggers
related to other artifacts that commonly occur in the signal at
this time (see Cycling-off Artifacts, FIG. 5). These are related to
acceleration pressure losses, which are difficult to compensate
for, or to phase delays between pressure and flow signals, which
are common in this setting, among other factors.
[0135] It should be pointed out that the selected values of K.sub.v
and K.sub.f may have little to do with actual patient elastance and
resistance. These values are simply used to facilitate detection of
phase transitions. As such the actual value of the signal does not
reliably reflect actual P.sub.mus, and such signals cannot be used
to reliably estimate the work of breathing or quantitative level of
pressure output by the patient.
[0136] FIG. 6 shows an example of signal Z generated from pressure,
flow and volume tracings. The signal was generated using a default
K.sub.f of 10, K.sub.f2 of 5.5 (ET tube #8) and a K.sub.v of 30.5
selected because it produced a flat baseline in the latter part of
expiration. Note the flat baseline of signal Z in the latter part
of expiration. In this patient, diaphragmatic electrical activity
was also monitored (lowest tracing), and this reflects the activity
of the main inspiratory muscle. Note the excellent agreement
between the onset of effort identified from the signal Z (arrows)
and the onset of diaphragm electrical activity. Note also that
T.sub.onset (arrows) was identified much earlier than the time at
which the ventilator triggered with a conventional triggering
algorithm (T.sub.trigger, top channel).
[0137] A number of approaches can be used to identify a change in
signal trajectory indicative of E.fwdarw.I transition
(T.sub.onset). Some of these include: [0138] a) Differentiating the
signal (.DELTA.signal/.DELTA.t) and comparing current values with
values obtained earlier. T.sub.onset is identified when the
difference exceeds a specified amount. [0139] b) Comparing current
values of signal with predicted values obtained from forward
projection of previous signal trajectory. T.sub.onset is identified
when the difference exceeds a specified amount. [0140] c) Comparing
current values of signal with values obtained earlier. T.sub.onset
is identified when the difference exceeds a specified amount.
[0141] d) Preferred approach: Differentiating the signal
(.DELTA.signal/.DELTA.t) and identifying points where
.DELTA.signal/.DELTA.t crosses zero in a positive direction
(t.sub.0(+)). The change in signal amplitude, relative to amplitude
at the immediately preceding t.sub.0(+), is continuously
calculated. T.sub.onset is identified when the difference between
current value and value at the preceding t.sub.0(+) exceeds a
specified amount (threshold). If the difference does not reach
threshold by the time .DELTA.signal/.DELTA.t crosses zero in a
negative direction (t.sub.0(-)), the difference is reset to zero,
until the next t.sub.0(+). This approach has the advantage of
filtering out slow, random undulations in baseline signal without
altering the relation between signal and inspiratory effort (which
would occur if a simple high pass filter were used). Such slow,
random undulations in baseline signal may be produced, for example,
by changes in thoracic blood volume, imperfect compensation for
mechanical non-linearities, or random changes in respiratory muscle
tone unrelated to phase transitions. The same approach can also be
used to estimate the amplitude of higher frequency baseline noise
(e.g. due to cardiac artifacts or secretions, see below). Such
information can then be used to automatically adjust the threshold
for identifying T.sub.onset.
[0142] Regardless of which approach is used to identify T.sub.onset
(a-d, above, or other approaches), a threshold must be set for the
magnitude of change that must be reached for T.sub.onset to be
declared. Several methods can be used to select such threshold.
Some of these include: [0143] i) A fixed threshold is arbitrarily
selected. For example, with approach (d), a signal increase, beyond
the latest t.sub.0(+), of 2 cmH.sub.2O may be used under all
conditions. Appropriate values may be chosen for other approaches.
Although feasible, when a universal threshold is used, the value
must be sufficiently high to avoid false auto-triggering under all
circumstances. Since noise level varies from patient to patient,
and from time to time, such a universal threshold would have to be
set to a level that is unnecessarily high under most conditions.
[0144] ii) Threshold may be individually selected by the user via
external controls. This can be achieved by the user selecting a
value that results in minimal auto-triggering. Alternatively, with
the help of graphical display of the signal, the user may adjust
the threshold above baseline noise level (e.g. horizontal dashed
line, FIG. 5). [0145] iii) Software algorithms can be developed to
distinguish noise from efforts and automatically adjust the
threshold accordingly.
[0146] The preceding account focussed primarily on identification
of E.fwdarw.I transitions. However, once K.sub.v and K.sub.f are
selected to produce a nearly flat baseline during expiration, the
shape of the signal during inspiration (but not necessarily its
amplitude, see above) provides a reasonable approximation of the
shape of inspiratory muscle output (P.sub.mus) (for example, see
FIG. 6). End of inspiratory effort (T.sub.end) is normally defined
as the point at which inspiratory muscle output rapidly declines
from its peak value. To implement this definition, the highest
value of signal Y (or Z) during the inflation phase can be
identified, in real time, using any of a number of standard
techniques. T.sub.end is identified when the signal decreases below
a specified value or a specified fraction of peak value.
[0147] At times, the signal undergoes a transient artifactual
reduction soon after ventilator triggering. An extreme example is
shown in FIG. 5 (arrow indicating Ventilator Trigger Artifact). It
is recognized as an artifact, as opposed to natural end of effort
(T.sub.end), because the signal resumes rising again. The presence
of these artifacts may cause false identification of T.sub.end. To
avoid this, if false T.sub.end identification occurs, the T.sub.end
identification circuitry is "blinded" to the signal for a set
period after T.sub.trigger (see T.sub.end Window Delay, FIG. 5) in
the same way the T.sub.onset identification circuitry is "blinded"
to the signal soon after ventilator cycling-off. Distinction
between artifactual and true T.sub.end can be easily made by the
consistent occurrence at T.sub.trigger and the secondary rise in
signal that characterize false T.sub.ends. The distinction can be
made by the user with the help of a monitor displaying the signal,
or by using software algorithms. The width of the T.sub.end Window
delay is adjusted accordingly.
[0148] At times, true T.sub.end occurs soon after ventilator
triggering. This is because inspiratory muscle activity can be
inhibited if inspiratory flow is high, and the ventilator
frequently delivers excessive flow soon after triggering. For this
reason, the procedure described above for T.sub.end identification
may, if used to cycle off the ventilator, result in medically
unacceptable inflation times. A back-up procedure is, therefore,
required to insure that the duration of inflation phase is
physiologically appropriate. The same procedure can be used to
insure that the inflation phase does not extend beyond
physiologically sound limits. The following is the rationale and
method for ensuring that the duration of the inflation phase
remains within physiologic limits.
[0149] In spontaneously breathing subjects and patients, the
duration of the inspiratory phase (T.sub.1) ranges between 25% and
50% of respiratory cycle duration (T.sub.TOT). In studies by the
inventor using proportional assist ventilation (PAV), with which
the duration of the ventilator's inflation phase mirrors the
patient's own T.sub.1, the ratio of T.sub.1 to T.sub.TOT
(T.sub.1/T.sub.TOT ratio) was also found to be between 0.25 and
0.5. Therefore, one approach to insure that the duration of the
inflation phase is within the physiologic range in modes in which
end of ventilator cycle is not automatically synchronized with the
patient is to constrain the duration of the inflation phase to be
between 0.25 and 0.5 of the total cycle duration of patient's own
efforts (to be distinguished from duration of ventilator cycles).
Accordingly, in another aspect of this invention, the end of the
ventilator cycle is constrained to occur within this physiological
range. Implementation of this procedure requires knowledge of the
true respiratory rate of the patient (as opposed to ventilator
rate). The true rate of the patient is the sum of ventilator rate,
the number of ineffective efforts occurring during the ventilator's
exhalation phase (arrows 1 to 3, FIG. 2) and the number of
additional efforts occurring during inflations triggered by an
earlier effort (arrows a to c, FIG. 2). The above-described method
for identifying T.sub.onset detects ineffective efforts occurring
during the ventilator's exhalation phase. These can be added to
ventilator rate. It may also be possible to identify extra efforts
occurring during the inflation phase of the ventilator (a, b, c,
FIG. 2) from the generated Y or Z signals. A simpler approach,
however, that is particularly suited for pressure support
ventilation, is to identify points in time at which flow begins
rising again during the inflation phase (FIG. 2). In pressure
support, flow typically declines progressively in the latter part
of inflation. The only possible explanation for a secondary rise in
flow, that is sustained for a significant duration (e.g. >0.3
second) is the occurrence of a second effort during inflation
(described by Giannouli et al, American Journal of Respiratory and
Critical Care Medicine, vol. 159, pages 1716-1725, 1999).
Identification of the extra efforts during the inflation or
exhalation phase can be made visually by the user (FIG. 2).
Alternatively, it can be done automatically using software or
analog circuitry. There are several possible approaches to
automatically obtain the number of extra efforts that did not
result in separate ventilator cycles. One such approach is to
differentiate the flow signal and determine the number of positive
and negative zero crossings of substantial duration (e.g. >0.4
second, to distinguish from high frequency noise and cardiac
artefacts). Another approach is to use Fourier frequency analysis
of the flow signal. There are clearly other mathematical approaches
to identify the characteristic flow transitions associated with
additional efforts. Thus, it is evident that there are many ways by
which true respiratory rate of patient can be determined.
[0150] Once the true respiratory rate of patient is known, it
becomes possible to calculate the real duration of respiratory
cycles of the patient (T.sub.TOT=60/respiratory rate) and determine
the range of inflation times consistent with a physiologic
T.sub.1/T.sub.TOT. For example, if patient's rate is 30/min,
T.sub.TOT is 2.0 seconds and the physiologic range for the
inflation phase is 0.5 to 1.0 second, reflecting a
T.sub.1/T.sub.TOT range of 0.25 to 0.50. Thus, according to this
aspect of the invention, average T.sub.TOT is determined using any
of a number of possible methods. The desirable duration of the
ventilator's inflation phase is then determined by multiplying
T.sub.TOT by a user selected physiologic T.sub.1/T.sub.TOT ratio or
a suitable default value (e.g. 0.4). In another implementation of
this method, a timer is reset at the onset of a new T.sub.onset or
a new ventilator cycle. The ventilator ignores other cycling-off
commands so long as time elapsed since the last T.sub.onset or
onset of ventilator cycle, is less than a set value (e.g. 0.3 of
T.sub.TOT). Similarly, to guard against excessively long ventilator
cycles, the timer may send a cycling-off command once time, since
the last T.sub.onset, or onset of inflation phase, exceeds a set
fraction of average T.sub.TOT (e.g. 0.45). The fractions used for
minimum and/or maximum cycling-off time can be fixed within the
ventilator or adjustable by the user.
[0151] An adaptation of this last aspect of the invention is
particularly suited for pressure support ventilation (PSV). Because
there is often some breath by breath variability in T.sub.TOT,
setting the end of ventilator cycle to a fixed fraction of average
T.sub.TOT results in some cycles having higher, and other cycles
having lower, T.sub.1/T.sub.TOT ratios. In this aspect of the
invention, only applicable to PSV, rather than causing the
ventilator to cycle-off at a predetermined time from the last
T.sub.onset, the ventilator is cycled off when inspiratory flow
reaches a specified amount, with this specified amount selected to
provide, on average, the specified T.sub.1/T.sub.TOT. This aspect
of the invention is implemented as follows: The interval between
successive inspiratory efforts (T.sub.TOT) is determined in several
elapsed ventilator cycles. The level of inspiratory flow at the
specified T.sub.1/T.sub.TOT fraction is noted. For example, if the
specified (desired) fraction is 0.4, and T.sub.TOT is 3.0 seconds,
flow is measured at 1.2 second after the preceding T.sub.onset
which triggered a ventilator cycle or, optionally, after the
trigger time of the relevant ventilator cycle. The average of
several such determinations, in several elapsed breaths, is used as
the cycling-off flow threshold in subsequent breaths. With this
approach, current cycles destined to have long T.sub.TOT
automatically have longer inflation cycles. This is so because
there is normally a correlation between the duration of inspiratory
muscle activity and the T.sub.TOT of individual breaths. Thus, in
breaths destined to have a long T.sub.TOT, inspiratory activity
tends to last longer and this, in PSV, delays the point at which a
specified cycling-off flow threshold is reached.
[0152] The information provided by the present invention can be
utilized in a number of ways: First, the time of T.sub.onset,
generated by the current invention, can be used to trigger
ventilator cycles by providing an appropriate signal to the
ventilator's triggering mechanism. Second, the end of the
ventilator inflation phase can be made to coincide with the end of
patient effort, as identified by the present invention, through
appropriate connections to the cycling-off mechanism of the
ventilation. Third, cycling-off can be made to occur at specified
times or, in the case of pressure support ventilation, at a
specified flow rate, after T.sub.onset or after the onset of
ventilator cycle. In this application, the user enters a desired
T.sub.1/T.sub.TOT ratio. The appropriate time, or flow, to
cycle-off is then determined from the inputted T.sub.1/T.sub.TOT
ratio and the value of average patient T.sub.TOT, obtained using
the present invention. Fourth, cycling off may occur at the
identified T.sub.end, conditional on this not violating a specified
minimum T.sub.1/T.sub.TOT ratio.
[0153] Whether or not it is used to synchronize the ventilator with
patient effort, the information provided by the current invention
can be displayed to the user to assist him/her in adjusting
ventilator settings to, indirectly, improve patient ventilator
interaction. In this connection, the information may be printed out
on command or be displayed on a monitor. The signal itself can be
displayed in real time along with other useful signals such as flow
and airway pressure. In addition, numerical values concerning
patient ventilator interaction can be displayed. Some recommended
values include: [0154] a) Trigger delay (difference between
ventilator trigger time and T.sub.onset). [0155] b) Cycling-off
error (difference between ventilator cycling-off time and end of
inspiratory effort). [0156] c) True respiratory rate of patient
(number of inspiratory efforts per minute). [0157] d) Average
duration between inspiratory efforts (T.sub.TOT). [0158] e) Number
of ineffective efforts, per minute or as a fraction of respiratory
rate. This is calculated as the difference between true rate of the
patient and ventilator rate. [0159] f) Number of central apneas (no
inspiratory efforts for a specified period, for example 10 seconds)
per hour, and/or % of time spent in central apnea. [0160] g) Flow
at a specified fraction of average T.sub.TOT in the pressure
support ventilation mode.
[0161] The numerical values may be accompanied by displayed
suggestions on how to adjust ventilator settings to reduce the
undesirable aspects of current interaction.
DESCRIPTION OF PREFERRED EMBODIMENT
[0162] The procedures of the present invention as described in
details above may be implemented in a device which may be
constructed as a freestanding device to be attached externally to a
ventilator, or may be incorporated within the ventilator. In either
case, the operation of the device requires inputs related to
pressure and airflow in the ventilator circuit. FIG. 7 shows a
design and components suitable for obtaining these signals.
Although it is possible to obtain these signals by attaching a flow
meter and pressure port to the common tube connecting ventilator to
patient 1, it is preferable to monitor flow and pressure separately
in the inspiratory and expiratory lines and to combine the signals.
This is to avoid clogging of the flow meter and to minimize the
number of tubing connections extending from near the patient's head
to the device. Accordingly, as shown in FIG. 7, a flow meter and
pressure port are inserted in the inspiratory line 2 and another
set is inserted in the expiratory line 3. Each set is connected to
appropriate pressure 4 and flow 5 transducers, which generate
signals proportional to pressure and flow, respectively. The
signals from each pressure 4 and flow 5 transducer is conditioned
with suitable low pass filters (e.g. 10 Hz) and offset and gain
circuitry. Suitable calibrations for the pressure and flow signals
are 10 cmH.sub.2O/volt and 1.0 l/sec/volt, respectively. The
processed inspiratory 6 and expiratory 7 flow signals are summed
using a summing amplifier 8 to produce a composite flow signal 9 to
be used by the device. The inspiratory 10 and expiratory 11
pressure signals are connected to a multiplexer 12. A comparator 13
receives the common flow signal 9 and provides a signal 14 to the
multiplexer 12 indicating the polarity of the flow signal 9. The
multiplexer generates a pressure signal 15 composed of the
inspiratory pressure signal 10 when flow is expiratory and the
expiratory pressure signal 11 when flow is inspiratory. In this
fashion the pressure 15 measured at any instant is a close
approximation of pressure in the tubing near the patient 1 since at
all times a static air column exists between the active transducer
and the common ventilator tubing 1 near the patient.
[0163] Pressure and flow signals are routinely generated in modern
ventilators using an approach similar to that of FIG. 7. If the
device of this invention is incorporated in the ventilator, the
pressure and flow signals generated independently by the ventilator
can be used instead.
[0164] FIG. 8 is a block diagram of an analog embodiment of the
invention. A summing amplifier 16 combines four signals, namely a)
the pressure signal 15 suitably inverted 17 and b) the flow signal
9 after suitable amplification 18 using a variable gain amplifier
19. This amplifier 19 provides the desired value of K.sub.f
(Equations 2, 3 and 4). c) A suitably conditioned and amplified
volume signal 20 generated by integrating 21 the flow signal (9)
after subtracting a highly filtered (using an ultra low pass filter
22) flow signal 23 to minimize volume drift. A variable gain
amplifier 24 provides the desired amplification (K.sub.v) of the
volume signal (as per Equations 3 and 4). d) A signal 25 comprised
of the product [flow*absolute flow] after suitable amplification.
This is an optional signal to be included if it is desired to
compensate for non-linearities in the pressure-flow relation as per
Equation 4. The signal corresponding to [flow absolute flow] is
generated by processing the flow signal 9 through an absolute value
circuit 26, and multiplying the output of this circuit 27 by the
flow signal 9 using an analog multiplier circuit 28. The resulting
signal 29 is then amplified with a variable gain amplifier 30 that
provides the desired value of K.sub.f2 (Equation 4).
[0165] The signal 31 generated by the summing amplifier 16 is
further processed by two circuits, one for detecting the onset of
inspiratory effort (T.sub.onset identification circuit 32) and one
for detecting the end of inspiratory effort (T.sub.end
identification circuit 33). The overall purpose of the first
circuit 32 is to measure the increase in the amplitude of the
signal 31 during periods in which the signal 31 is rising, within a
specific time window in the breath determined by a T.sub.onset
window circuit 34. This time window begins after a specified delay
35 from the point at which expiratory flow decreases below a
specified value (e.g. -0.2 l/sec) during expiration. As seen in the
diagram of the first circuit 32, the signal 31 is differentiated
using a differentiator 36. The differentiated signal 37 is filtered
using an appropriate low pass filter (e.g. 5 Hz) 38 to remove high
frequency noise. The filtered differentiated signal 39 is passed
through two comparators. One comparator 40 sends an enabling
positive signal 41 when the filtered differentiated signal 39 is
positive and the other comparator 42 sends an enabling positive
signal 43 when the filtered differentiated signal 39 is negative.
The unfiltered signal 37 is integrated 44 when two gates 45,46 are
enabled. The first gate 45 is enabled when the filtered
differentiated signal 39 is positive, as detected by the positive
comparator 40. The second gate 46 is enabled during the specified
time window during expiration, as detected by the T.sub.onset
window circuit 34 and conveyed to the gate by an enabling signal
47. The integrator 44 is reset whenever the filtered differentiated
signal 39 becomes negative as detected by the negative comparator
42. In this fashion integration begins anew only when the signal is
rising within the specified time window. The integrator output 48
is received by a comparator 49 which sends out a signal 50,
indicating T.sub.onset, when integrator output exceeds a specified
threshold set by an external EI threshold adjust 51.
[0166] The specific design used for detection of onset of effort in
this implementation 32 is selected because it offered an optimal
combination of sensitivity and specificity (i.e. sensitive yet not
prone to false triggering). It is clear, however, that other
designs for detecting a change in signal trajectory are possible.
For example, the filtered differentiated signal 39, representing
current rate of change in signal, can be delayed by a specified
amount (e.g. 200 msec). A comparator (not shown) compares the
current and delayed forms of the filtered differentiated signal. A
signal, indicating onset of effort, is generated when the
difference exceeds a threshold value. Alternatively, the signal
itself 31 may be delayed by a specified amount (e.g. 200 msec). A
comparator (not shown) compares the current and delayed forms of
the actual signal and generates a signal, indicating onset of
effort, when the difference exceeds a threshold value. Other
approaches are possible within the scope of this invention.
[0167] For identifying the T.sub.end 33, the signal 31 is first
differentiated 52 and the differentiated signal 53 is reintegrated
54. The integrator is reset at the onset of inspiratory effort
(T.sub.onset) using the signal 50 generated from the T.sub.onset
identification circuit 32. In this fashion, any baseline offset in
the signal 31 is eliminated and the output of the integrator 55
reflects only the increase in signal 31 amplitude from T.sub.onset.
Integrator output 55 is connected to a peak detector circuit 56,
which is also reset by the T.sub.onset signal 50. The output of the
peak detector 57 is attenuated 58 with a suitable attenuation
factor (e.g. 50%). Optionally, the attenuation factor may be
individually adjusted by the user through an external input 59. A
comparator 60 sends a signal 62 when current integrator output 55
decreases below the attenuated peak detector output 61. In this
fashion the end of inspiratory effort is detected when the current
integrator output 55 decreases below a set percent of the peak
level reached during the current inspiratory effort.
[0168] At times, the signal 31 or 55 transiently decreases at the
time of ventilator triggering (Ventilator Trigger Artifact, FIG.
5). Unless corrected, or allowed for, this artefact may result in
false detection of T.sub.end. A circuit is incorporated to reduce
or eliminate the occurrence of false identification of T.sub.end.
The circuit consists of a delay circuit 63 similar to the one used
in the T.sub.onset identification circuit 34. A timer is activated
by a T.sub.trigger signal 64 received from a T.sub.trigger
identification circuit 65. The latter circuit receives inputs from
the pressure 15 and flow 9 signals. The pressure signal is
differentiated 66 and the resulting signal 67 is directed to a
comparator 68 with a suitable reference value (e.g. 15
cmH.sub.2O/sec). The flow signal 9 also is connected to a
comparator 69 with a suitable reference value (e.g. 0.3 l/sec). The
outputs of the two comparators 68,69 are received by an OR gate 70
which sends a T.sub.trigger signal 64 to the delay circuit 63 when
either the differentiated pressure signal 67 or the flow signal 9
exceed the set value in the respective comparator 68 or 69. The
delay circuit 63 in turns sends a signal 71 to an AND gate 72 after
a specified delay set either externally via a user input 73 or
internally as a default value (e.g. 0.2 sec). The AND gate 71 also
receives the T.sub.end signal 62 and sends a final T.sub.end signal
74 only if it occurs after the specified delay from T.sub.trigger.
In this fashion, T.sub.end signals generated by the triggering
artifacts are screened out.
[0169] User Inputs:
[0170] The number and types of user inputs may vary depending on
how comprehensive the device is and the extent to which user
involvement is desired by the manufacturer. In the most
comprehensive analog embodiment shown in FIG. 8, there are seven
user inputs: [0171] 1) K.sub.f adjust 75: This input determines the
gain of the K.sub.f variable gain amplifier 19. A suitable range is
1 to 25 cmH.sub.2O/l/sec. Because the calibration factors of the
flow and pressure signals may be different (for optimal signal to
noise ratio, see above) an attenuation factor is incorporated to
make allowance for the different calibration factors. For example,
if the flow calibration factor is 1.0 l/sec/volt and the pressure
calibration factor is 10.0 cmH.sub.2O/volt, the relation between
the K.sub.f adjust input 75 and the gain of the K.sub.f variable
gain amplifier 19 should be 10. In this fashion, the output of the
K.sub.f variable gain amplifier, which has units of pressure, is
comparable to the pressure signal 17 at 10.0 cmH.sub.2O/volt.
[0172] 2) K.sub.f2 adjust 76: This input determines the gain of the
K.sub.f2 variable gain amplifier 30. A suitable range is 1 to 25
cmH.sub.2O/l.sup.2/sec.sup.2 to take account of the various sizes
of endotracheal tubes used in practice. Again, a suitable
attenuation factor between the K.sub.f2 input 76 and the actual
K.sub.f2 gain 30 needs to be incorporated to allow for differences
in pressure and flow calibration factors (see #1 immediately
above). [0173] 3) K.sub.v adjust 77: This input determines the gain
of the K.sub.v variable gain amplifier 24. A suitable range is 5 to
100 cmH.sub.2O/l. A suitable attenuation factor between the K.sub.v
input 77 and the actual K.sub.v gain 24 needs to be incorporated to
allow for differences in pressure and flow calibration factors (see
#1 immediately above). [0174] 4) T.sub.onset window delay 35: This
input determines the desired delay, from the point at which
expiratory flow decreases below a set value, before the device
begins looking for T.sub.onset. A suitable range is 0 to 3.0
seconds. [0175] 5) E I threshold 51: This determines the amount of
increase in signal amplitude, as detected by the integrator 44 of
the T.sub.onset circuit 32, above which T.sub.onset is identified.
A suitable range is 0.1 to 10.0 cmH.sub.2O. [0176] 6) Signal
attenuation factor 59: This determines how much signal amplitude
must decrease, after T.sub.onset, before the T.sub.end is
identified. A suitable range is 20 to 90%. [0177] 7) T.sub.end
Window delay 73: This input determines the period, from
T.sub.trigger, during which T.sub.end signals 62 are screened out.
A suitable range is 0.0 to 0.3 second.
[0178] Some inputs may be deleted by using fixed default values
within the device. For example, the K.sub.f adjust input 75 may be
deleted and a fixed value of 10.0 is used. A fixed T.sub.onset
delay value of, for example, 0.3 second may be used, eliminating
the T.sub.onset window delay input 35. A suitable default signal
attenuation value (e.g. 50%) may be used replacing the
corresponding input 59. Likewise, a T.sub.end window delay of 0.2
second may be used eliminating the T.sub.end window delay 73.
Clearly, the more fixed the settings are the less reliable the
performance of the device may become. However, this may be
acceptable under some circumstances with the potential benefit of
simplifying the operation of the device. An alternative would be to
have the device operate with default settings unless changed by the
user.
[0179] Other inputs may also become unnecessary if simpler forms of
the signal 31 are generated. For example, signal component related
to the non-linear flow function 25 may be eliminated according to
Equation 3. In this case the K.sub.f2 adjust input 76 is deleted.
Likewise, the signal component related to volume 20 may be
eliminated, according to Equation 2, with corresponding deletion of
the K.sub.v adjust input 77. Again, the simpler the device, the
less reliable its performance will become but this may be
acceptable under certain circumstances. In its simplest form, all
the user needs to do is to set the E-I threshold input 51.
[0180] Device Outputs:
[0181] Certain internal signals need to be displayed to allow the
user to adjust the input settings, while others provide the user
with the results of monitoring. These signals can be displayed on a
monitor 78 included in a freestanding device. Alternatively, if the
device is incorporated inside the ventilator, the monitor of the
ventilator can be used for this purpose. A third embodiment
involves directing the device's outputs to an analog to digital
converter and displaying the outputs on a separate computer.
[0182] The following output signals are necessary for adjusting the
input settings:
[0183] a) The main signal itself 31.
[0184] b) The output of the integrator 48 in the T.sub.onset
circuit (32).
[0185] The use of these two signals for the sake of input
adjustment is described below under OPERATION (below).
[0186] Additionally, the signals representing flow 9, pressure 15
and volume 79 may be displayed on the monitor for general
monitoring purposes.
[0187] Signals representing the onset of inspiratory effort 50
(T.sub.onset) and end of inspiratory effort 74 (T.sub.end) are also
displayed on the monitor. In the event these signals are to be used
to actively control the cycling of the ventilator, they are
communicated to the ventilator's cycling mechanism.
[0188] Additional information of value in guiding ventilator
setting is most conveniently generated by a small microprocessor. A
block diagram of a preferred embodiment (103) is provided in FIG.
9. Here, the flow signal 9 is digitized using an analog to digital
converter. In addition, the central processing unit receives the
signals corresponding to T.sub.onset 50 and T.sub.end 74 of
inspiratory effort and signals reflecting ventilator trigger time
(T.sub.trigger) 64 and cycling off time (T.sub.off, derived from
the T.sub.onset Window circuit 34 see also 96 in FIG. 12). The
latter two signals may also be obtained directly from the
ventilator. The user inputs the ventilator mode 88 and the desired
T.sub.1/T.sub.TOT ratio 89. From these data, the microprocessor
calculates trigger delay (T.sub.trigger-T.sub.onset) 80 and cycling
off delay (T.sub.off-T.sub.end) 81. The flow signal is
differentiated. Additional inspiratory efforts during the inflation
phase 82 are identified when the differentiated flow becomes
positive after an earlier negative phase (Identify additional
efforts function 82, FIG. 9). A "calculate patient rate function"
(83, FIG. 9) calculates respiratory rate of patient 83 from the sum
of number of T.sub.onset transitions during expiration 50 in the
last minute and the number of additional efforts during inflation
82 in the last minute. The number of ventilator cycles per minute
84 is calculated from the number of T.sub.trigger signals 64 in the
last minute. The number of ineffective efforts 85 is calculated
from the difference between patient respiratory rate 83 and
ventilator rate 84. This additional information is then displayed
on the monitor. Additionally, with knowledge of patient respiratory
rate 83 the average breathing cycle duration (T.sub.tot,
T.sub.tot=60/respiratory rate) of the patient can be calculated.
The microprocessor calculates the desirable duration of ventilator
cycle 87 (desirable T.sub.1=T.sub.TOT*desirable T.sub.1/T.sub.TOT)
where T.sub.1/T.sub.TOT is a default value (e.g. 0.4) or a user
input 89. The microprocessor also receives a user input indicating
the mode of ventilation 88. In the PSV mode, the microprocessor
samples flow at the desired T.sub.1 in several elapsed cycles (Flow
at desired T.sub.1 function 90), and displays the average value on
the monitor. The user can take advantage of this information
(desired T.sub.1 or Flow at desired T.sub.1) to adjust ventilator
settings to result in optimal T.sub.1/T.sub.TOT.
[0189] In another embodiment of the output processor 103 patient
respiratory rate (or T.sub.TOT) is inputted to the processor,
replacing the "Calculate Patient Rate" function 83. This input is
then used to calculate the "Desirable T.sub.1" 87 and "Flow at
Desired T.sub.1" 90. Patient respiratory rate may be determined by
the user from inspection of chest movements or by observing the
flow tracing on the monitor, or automatically using computational
methods other than the ones described in the above embodiment
103.
Operation
[0190] When the device is built inside the ventilator, the pressure
15 and flow 9 signals are permanently connected to the device. For
freestanding systems, the first step is to connect the flow meters
and pressure ports to the inspiratory 2 and expiratory 3 lines
close to the ventilator (FIG. 7). The device is turned on. Tracing
of the Signal 31 appears on the screen (FIG. 10). Subsequent steps
depend on what inputs are available on the device and user
preference. For the most comprehensive analog embodiment (FIG. 8)
the recommended procedure is as follows: [0191] 1) Enter the
K.sub.f2 value 76. This is the K.sub.2 value of the endotracheal
tube in use. A table is provided that states the K.sub.2 values for
the range of endotracheal tube sizes used. [0192] 2) Set the other
inputs to default values as follows: K.sub.f 75=10; K.sub.v 77=25;
T.sub.onset window delay=10% of respiratory cycle duration. For
example if respiratory rate is 20/min, set the delay to 0.3 sec;
Signal attenuation factor 59=50%; T.sub.end Window Delay 73=0.2
second. [0193] 3) If the baseline of the signal 31 is flat in the
latter half of expiration (e.g. 91, FIG. 10B) no further adjustment
of K.sub.v is necessary. If it is not flat (e.g. 92, FIG. 10A),
adjust K.sub.v setting 77 to make it flat or slightly sloping down
(e.g. 91, FIG. 10B). [0194] 4) If it is difficult to have
reasonably linear signal trajectory using the K.sub.v adjust input
77 alone, adjust the K.sub.f2 adjust input 76 up or down as
necessary to minimize non-linearities. [0195] 5) The tracing
representing integrator output 48 (FIG. 10B) shows relatively large
broad waves, representing inspiratory efforts 93, FIG. 10B), and
smaller, briefer spikes representing noise (94, FIG. 10B). Set the
E I threshold level (51) to be just above the smaller spikes in
several consecutive breaths (e.g. 95, FIG. 10B). [0196] 6) Display
the T.sub.onset 50 and T.sub.end 74 on the screen. If there are
frequent T.sub.onset signals triggered by noise, increase the level
of the E I threshold. If there are frequent T.sub.onset signals
triggered early in expiration, increase the T.sub.onset Window
delay 35. If the T.sub.end signal occurs too early or too late
during the declining phase of the signal 31, adjust the signal
attenuation factor 59 accordingly. If there are frequent false
triggers of T.sub.end at the time of ventilator triggering,
increase the T.sub.end Window Delay 73 to eliminate false
triggering of T.sub.end.
[0197] FIGS. 11 and 12 show details of the electrical circuitry
used in the preferred embodiment (FIG. 8). All circuits are powered
by a suitable +/-8 volts power supply. In FIG. 11 the circuitry
used to generate the main signal 31 (block 80 in FIG. 8) is
displayed. The summing amplifier 16 with its 4 inputs (17,18,20,25)
is shown in the lower right corner of the Figure. Circuitry used to
process the four inputs prior to the summing amplifier stage is
outlined in boxes bearing the same numbers as in the corresponding
components of the block diagram (FIG. 8). The individual electrical
components in each circuit are identified by standard electrical
symbols and the values of resistors and capacitors indicated are
those used in a properly functioning prototype.
[0198] FIG. 12 shows details of the electrical circuitry used for
T.sub.onset identification 32 and T.sub.end identification 33. As
in FIG. 11, the individual electrical components in each circuit
are identified by standard electrical symbols and the values of
resistors and capacitors indicated are those used in a properly
functioning prototype. The specific function of each circuit and
its connections to other circuits have been described in detail in
relation to the block diagram of FIG. 8, and the design of each
circuit is standard for the purpose intended in each case. Some of
the component circuits need additional explanation, however:
[0199] T.sub.onset Window circuit 34: In this circuit the flow
signal is connected to a Schmitt trigger circuit (left half of the
T.sub.onset window circuit 34) characterized by hysteresis. With
the indicated values of the different circuit components, the
Schmitt circuit sends out a constant voltage (8 volts) whenever
flow decreases below -0.2 l/sec 96. The signal 96 remains on
until-flow rises to >0.2 l/sec. In this application, the onset
of the signal 96 indicates the beginning of the exhalation phase
and is also used to mark the end of ventilator cycle (T.sub.off).
The output of the Schmitt trigger circuit is connected to a delay
circuit with an externally adjustable delay time 35. The output of
the delay circuit 97 is received by an AND gate 98. The AND gate 98
also receives the output of the Schmitt trigger circuit 96 directly
and sends a signal when the T.sub.onset window is open, as
indicated by the output of the Schmitt trigger circuit 96 but only
after the specified delay 35 has elapsed, as indicated by a
positive output from the delay circuit 97. In turn, the output of
the AND gate 47 (referred to as Q signal in FIG. 12) serves
multiple functions that include enabling one of the transmission
gates 46 in the T.sub.onset identification circuit 32.
[0200] T.sub.trigger detection circuitry: There are many ways by
which the time at which the ventilator was triggered can be
detected. In this embodiment T.sub.trigger was detected when the
rate of increase in pressure exceeded 15 cmH.sub.2O/second OR flow
increased beyond 0.4 l/second. To this end, a differentiator 66 was
used to obtain .DELTA.pressure/.DELTA.t 67. Next, a comparator 68
produces a positive signal 99 when .DELTA.pressure/.DELTA.t 67
exceeds a set value of 15 cmH.sub.2O/second. In another circuit 69
a comparator generates a positive signal 100 when flow (9) exceeds
0.4 l/second. Two diodes 101,102 function as an OR gate so that a
positive signal (T.sub.trigger, 64) is generated when either the
.DELTA.pressure/.DELTA.t or flow exceed the set respective
thresholds.
[0201] T.sub.end Window circuit 63: This circuit has four
components. First, the Q signal 47, representing time window for
T.sub.onset detection, is inverted using an inverter 104. The
positive phase of this inverted Q signal 105 defines the maximum
period during which T.sub.end can be located. The second component
is an AND gate 106 which receives the inverted Q signal 105 and the
T.sub.trigger signal 64 and sends a positive signal 107 when both
its inputs are positive. The positive edge of this signal 107
activates a flip-flop switch 108, which is the third component of
the T.sub.end Window circuit. The flip-flop switch 108 is reset by
the Q signal 47. The fourth component is a delay circuit 110 with
an adjustable external control 111. The delay circuit 110 receives
the output of the flip-flop switch 109. After the set delay, the
delay circuit 110 sends out a positive signal 112, which persists
until the beginning of the Q signal 47. The output of the delay
circuit 112 is one of the two inputs to the main AND gate 72 which
generates the T.sub.end signal 74.
[0202] The other components of the T.sub.end circuit 33, as shown
in FIG. 12, include the differentiator 52 and integrator 54 that
calculate the change in the main signal 31 since the onset of the
current effort (55, referred to as S' in FIG. 12). The integrator
is reset by the T.sub.onset signal 50. The peak detector that
determines the highest level of signal S' 55 reached during the
current effort, is shown 56 and is also reset by the T.sub.onset
signal 50. The output of the peak detector 57 is attenuated with an
externally adjustable attenuator 58. Finally, a comparator 60
receives the output of the attenuated peak signal 61 and the
differentiated integrated signal (55, S') and sends a T.sub.end
signal 62 when the latter 55 decreases below the former 61. The
T.sub.end signal 62 is gated out only if the T.sub.end Window is
open as indicated by a positive output of the T.sub.end Window
delay circuit 112. This gating function is performed by an AND gate
72.
[0203] The circuitry used in this preferred embodiment is clearly
not the only way by which the functions and results contemplated by
the current invention can be implemented. Other circuit designs can
be used to accomplish the same objectives and these are within the
scope of this invention.
[0204] FIGS. 13 to 17 show flowcharts for the different functions
performed by the output microprocessor (FIG. 9). The power on
start-up routine (113, FIG. 13) clears the memory and enables the
Interrupt Request (IRQ) Process. The IRQ process (114, FIG. 14) is
executed at suitable intervals (e.g. every 5 msec). It collects
data from various inputs (see FIG. 9 for inputs), calculates the
time derivative of flow and stores collected and derived data in
memory. It also checks for the times at which T.sub.trigger,
T.sub.onset, T.sub.off, and T.sub.end occur and stores these times
in memory. Because all these timing inputs are square functions,
detection of the times at which these events occurred is based on a
simple comparison of current value with the immediately preceding
one. If current value is high and preceding value is low, the event
is deemed to have occurred. For example, if current value of
T.sub.trigger, input is high while the immediately preceding value
was low, T.sub.trigger is deemed to have occurred then, and so on.
Finally, the IRQ process writes the waveform data and calculated
variables to the monitor. The main program loop (115, FIG. 15)
performs the various functions identified in the block diagram
(FIG. 9) in sequence each time a T.sub.trigger is detected. The
flow charts of individual functions are shown in individual
diagrams bearing the same numbers. In the trigger delay function
(80, FIG. 15), when the difference between current T.sub.trigger
and the last T.sub.onset is >1.0 second, trigger delay is
ignored. Thus, the maximum trigger delay allowed is 1.0 second.
Situations in which T.sub.onset occurs more than 1.0 second before
T.sub.trigger are usually ventilator cycles triggered by the
ventilator and not by the patient. The Calculate Ventilator
T.sub.TOT function (117, FIG. 15) calculates the interval between
current and previous T.sub.trigger. The cycling off delay function
(81, FIG. 16) calculates the difference between the end of
ventilator cycle (T.sub.off) and end of patient effort ((T.sub.end)
in current cycle. In the Identify Additional Efforts function (82,
FIG. 17) the program looks in the interval between T.sub.trigger
and T.sub.off of the previous cycle for points at which
.DELTA.flow/.DELTA.t crosses from negative to positive and stays
positive for 300 msec. When this occurs, it identifies an
additional effort during the previous ventilator cycle and adds its
time to the circular buffer for subsequent counting. The choice of
300 msec is quite conservative and may suitably be reduced to 200
msec or even less. In the Calculate Patient Rate function (83, FIG.
17) the program calculates the number of T.sub.onset transitions
and number of Additional Efforts during inflation in the one-minute
interval before the current T.sub.trigger. In this chart PTE refers
to efforts occurring during the exhalation phase of the ventilator
(i.e. T.sub.onset transitions) and PTAE refers to Additional
Efforts occurring during the inflation phase of the ventilator. In
the Desirable T.sub.1 calculate function (87, FIG. 16) the average
patient cycle duration (T.sub.TOT) is calculated from 60/patient
respiratory rate calculated in the preceding function (83).
Desirable T.sub.1 is then calculated from average patient T.sub.TOT
and the desirable T.sub.1/T.sub.TOT as indicated by the desired
T.sub.1/T.sub.TOT input (89, FIG. 9). In the Calculate Target Flow
function (90, FIG. 16) the first decision is whether the mode is
pressure support ventilation (PSV). If so, the program reads flow
at an appropriate time in the immediately preceding ventilator
cycle. There are a number of options for the appropriate time at
which to measure flow (see next paragraph). Occasionally, the time
at which to measure flow may occur after the end of the ventilator
cycle, where flow is negative (i.e. expiratory). This is the case
when the respiratory time constant of the patient
(resistance/elastance) is too short. A provision is made whereby if
flow at the chosen time is negative it is assigned a value of zero.
With certain variables it is preferable to provide the user with
average results as opposed to, or in addition to, results of
individual cycles, which may be quite variable. For this reason
individual results of certain variables are stored in circular
buffer (e.g. Trigger delay (80, FIG. 15), Cycling off delay (81,
FIG. 16), Ventilator T.sub.TOT (117, FIG. 15) and Target flow for
end of cycle (90, FIG. 16)). The Calculate Averages function (116,
FIG. 16) then calculates the average values in a preset number of
elapsed breaths. In the illustrated embodiment (116, FIG. 16), the
number of cycles averaged is 10. However, other numbers may be
chosen depending on manufacturer or user preference. Two other
variables are derived from these averaged values. Ventilator rate
(84, FIG. 15) is calculated from [60/average ventilator T.sub.TOT
(117)) and the number of Ineffective Efforts (85, FIG. 16) is
calculated from the difference between Average Patient Rate (83,
FIG. 17) and Average Ventilator Rate (84, FIG. 15).
[0205] In the illustrated embodiment for calculating target flow to
cycle off pressure support ventilation (90, FIG. 16) the point
chosen to measure flow was the preceding T.sub.trigger plus an
interval corresponding to desirable T.sub.1, with the latter based
on desired T.sub.1/T.sub.TOT and average respiratory cycle of
patient efforts (87, FIG. 16). There are, however, several other
options for selecting the point in time at which to measure flow.
These include, but are not limited to: [0206] a. Desirable T.sub.1
is added to the T.sub.onset preceding the previous T.sub.trigger
instead of adding it to T.sub.trigger itself. [0207] b. Desirable
T.sub.1 is added to a point in time between previous T.sub.trigger
and the preceding T.sub.onset. [0208] c. Desirable T.sub.1 is
calculated from desired T.sub.1/T.sub.TOT fraction of the T.sub.TOT
of the individual patient cycle that included the previous
T.sub.trigger. This value is then added to the previous
T.sub.trigger, the preceding T.sub.onset or some intermediate
time.
[0209] Each of these options has advantages and disadvantages. In
practice, the difference in net result should be small. However,
some manufacturers or users may prefer one or the other or even a
completely different option.
[0210] The resulting output of such microprocessor (FIG. 9) are
displayed on a monitor. The information provided can be utilized by
the user to adjust ventilator settings in order to optimize
patient-ventilator interaction. Alternatively, or in addition, some
of the outputs can be channeled to the cycling mechanism of the
ventilator to effect such optimization automatically. Of particular
utility is the use of information provided by the Desirable T.sub.1
function 87 to automatically set the duration of the inflation
phase of the ventilator. Likewise, the output of the Target. Flow
for End of Cycle in the pressure support mode 90 can be used to
automatically determine the flow threshold at which the ventilator
cycles off in this mode. Other examples of use of generated data
include, but are not limited to, increasing the flow threshold for
cycling off pressure support when the Cycling off Delay function 81
produces large positive values or when the Calculate Ineffective
Efforts function 85 indicates a large number or fraction of such
efforts. The magnitude of pressure support (i.e. amount of increase
in pressure at triggering) may also be automatically decreased in
the presence of long trigger delays, as unveiled by the Calculate
Trigger Delay function 80, long and positive Cycling off Delays
(per 81) or excessive ineffective efforts (per 85). Microprocessor
output can thus be used for closed loop control of amplitude and
duration of ventilator assist.
[0211] Whereas the preferred embodiment described herein utilized
electrical circuitry to generate the Signal and to determine
T.sub.onset and T.sub.end, it is clear that any and all the
functions executed by electrical circuitry for the current
application can be readily executed by digital technology. FIG. 18
is a block diagram illustrating one embodiment of a fully digital
device. The device receives the various inputs either via an A/D
converter or directly to the central processing unit (CPU)
depending on whether the primary inputs are in digital or analog
form. In its most comprehensive form, these inputs include pressure
15, flow 9, K.sub.f 75, K.sub.f2 76, K.sub.v 77, T.sub.onset window
delay 35, T.sub.end window delay 73, Signal attenuation factor 59,
EI threshold 51, Mode 88 and desired T.sub.1/T.sub.TOT 89. The
microprocessor executes some functions in real time and others in
non real time when a T.sub.trigger is identified. The non real time
functions are similar to those described in detail in relation to
the output microprocessor of FIG. 9 and the associated flow charts
of FIGS. 13 to 17. These will not be described further. The real
time functions are executed at suitable intervals; every 5 to 10
msec being optimal. The timed IRQ process 118 is illustrated in
flow chart form in FIG. 19. After reading and storing the various
inputs, it calculates volume and flow.sup.2. The rate of change in
pressure is calculated for use in the Calculate T.sub.trigger
function 119 and the rate of change in flow is calculated for use
in the Identify Additional Efforts function 82. The main Signal is
then calculated according to Equation 4 and Signal is
differentiated for use in the T.sub.onset and T.sub.end
identification functions 121,123. T.sub.trigger is then looked for
using a T.sub.trigger calculate function 119 and, if found, the
T.sub.rigger flag is set to TRUE. This initiates the non real time
functions. Subsequently, the T.sub.onset Window calculate function
120 is used to determine whether this window is open and, if so,
the Calculate T.sub.onset function 121 is processed to determine
whether a T.sub.onset occurred. Finally, the Calculate T.sub.end.
Window function 122 and the Calculate T.sub.end function 123 are
processed to identify if a T.sub.end occurred. The T.sub.trigger
calculate function 119, T.sub.onset Window calculate function 120,
Calculate T.sub.onset function 121, Calculate T.sub.end Window
function 122, and the Calculate T.sub.end function 123 are
illustrated in flow chart format in FIGS. 19 to 21. These charts
are self-explanatory particularly in light of the detailed
description of the same functions in relation to the block diagram
(FIG. 8) and circuit diagrams (FIGS. 11 and 12) of the analog
implementation.
[0212] As in the case of the analog implementation, the digital
implementation can be simplified to different degrees depending on
user and manufacturer preferences. The outputs of the device may
also be expanded or reduced to meet user needs.
SUMMARY OF DISCLOSURE
[0213] In summary of this disclosure, the present invention
provides a method and apparatus for detecting the onset and the end
of inspiratory effort in a patient on mechanical ventilation.
Modifications are possible within the scope of the invention.
* * * * *