Apparatus for the monitoring and control of respiration

A respiration monitor for measuring the respiration of a subject on a mechanical ventilator comprises first transducer means arranged to produce a first signal representative of a respiration status of the subject, second transducer means arranged to produce a second signal representative of a cyclical status of the mechanical ventilator, and processor means arranged to receive said first and second signals and to produce an output signal dependent upon the relative phase of said first and second signals. The invention extends to a method of monitoring, and also to a method of determining the rate at which a mechanical ventilator should be set in order to achieve 1:1 entrainment with spontaneous respiratory efforts.

FIELD OF THE INVENTION 
The present invention relates to a method and apparatus for the monitoring 
and control of a subject's physiological status (e.g. respiration). It 
finds particular (although not exclusive) application in monitoring and 
controlling chaotic activity during ventilation of newly born infants, and 
in the provision of a visual display which is suitable for diagnostic 
purposes. 
BACKGROUND OF THE INVENTION 
Patients with severe respiratory illness may require assistance with their 
breathing if their lungs are stiff, the respiratory muscles weak, or if 
oxygenation of the blood is inadequate due to lung disease. Respiratory 
assistance is given by blowing air/oxygen mixtures into the lungs using a 
mechanical ventilator to expand the lungs and take over some, or all, of 
the work of breathing. 
Patients who are often breathing spontaneously during mechanical 
ventilation may "fight the ventilator" when they are trying to breathe at 
different times from the action of the ventilator. This creates problems 
in the exchange of gases in the lungs, can lead to sudden changes in the 
action of the heart and in the blood pressure, and can affect the flow of 
blood to the brain when the patients are severely ill. These adverse 
effects of "fighting the ventilator" are seen most dramatically in the 
sick, prematurely born infant. These tiny infants are at risk from brain 
damage when their breathing patterns become disordered, and more efficient 
methods of mechanical ventilation are constantly being sought for this 
group. Rapid changes in clinical state, irritability and the presence of 
airway reflexes lead to complex and rapidly changing interactions between 
the baby and the mechanical ventilator. Muscle paralysis is used to 
suppress spontaneous respiratory efforts, but may be associated with 
cardiovascular compromise and the need for higher inflating pressures. New 
techniques of mechanical ventilation attempt to induce and maintain 
phase-locking or phase-synchrony by the use of fast rates and short 
inspiratory times, or by triggering ventilator inflation using sensors to 
detect diaphragmatic excursion. In clinical practice, mechanical 
ventilation of the newborn is hampered by an inability to assess 
baby-ventilator interactions by clinical observation at the high 
spontaneous respiratory frequencies seen (typically 1-2 Hz). Spontaneous 
respiratory activity is frequently erratic and accompanied by 
unpredictable activity such as hiccoughs, gasps and responses to painful 
and other stimuli. An ideal respiratory monitoring system would be able to 
track the phase, amplitude and frequency of spontaneous respiratory 
activity relative to mechanical inflation from breath to breath. 
The physiological interactions underlying cardiorespiratory control are 
usually non-linear in nature. Entrainment of biological oscillatory 
rhythms, such as spontaneous respiratory activity, can be achieved under 
certain conditions by the application of a periodic stimulus, such as 
mechanical ventilation, provided that sufficient afferent information 
reaches the rhythm generator to bring about entrainment. During stable 
entrainment the output frequency of the spontaneous oscillator will be 
drawn into simple integer relationships with that of the periodic 
stimulus, and a fixed phase relationship will be maintained indefinitely, 
provided stochastic noise is minimal. 
In order to achieve an adequate description of the complex changes which 
characterise the response of stimulated nonlinear systems in physiology we 
have developed a method (the frequency tracking locus) of tracking 
cycle-by-cycle changes as opposed to the steady state response. The 
advantage of the frequency tracking locus method is that it allows a 
quantitative estimate of the state of entrainment in a stimulated system, 
as well as providing a visualisation of the interactions between the 
stimulus and the output from the system. 
The frequency response of a linear system may be determined by applying 
sinusoids of a fixed frequency and calculating the amplitude and phase 
difference between input and output. The steady-state frequency response 
of the system can be determined over its entire range by applying input 
sinusoids incrementally. In nonlinear oscillatory systems the input and 
output do not exhibit either a fixed amplitude ratio or phase 
relationship. 
A key factor in the analysis of the interaction of nonlinear oscillations 
is the ability to track frequencies and transients in both stimulus and 
output signals. Under steady state conditions frequency tracking can be 
achieved by the use of Fourier estimators, but as entrainment of a 
nonlinear oscillator becomes unstable, the output oscillation becomes 
non-stationary. The Fourier integral is based on the assumption that the 
data extend over an infinite range without any change in frequency content 
i.e. the waveform is stationary. In practice this condition is never met, 
but accurate, practical frequency resolution can be achieved with a 
minimum of approximately 3 to 5 cycles of the fundamental frequency (the 
lowest frequency in the waveform). Consequently, Fourier estimation has 
been successful in those biological studies where experimental design has 
determined the stationarity of the frequency content. Examples of such 
studies include monitoring thermal entrainment of physiological rhythms, 
heart rate variability, controlled breathing experiments and observations 
of the effects of respiration on blood pressure in the newborn. When 
studies of physiological systems involve the analysis of spontaneous 
activity, however, it is known that nonlinearities in the control 
structure induce non-stationarities in the associated waveforms. Hence, 
Fourier estimators are unable to track the shorter periods of stationarity 
which occur. In this case we and others have applied linear estimation 
methods which can resolve over stationary data lengths of 1.5 cycles of 
the fundamental. While these methods have proved useful, they have two 
principal disadvantages in relation to the study of nonlinear 
oscillations. First, a great deal of care must be taken to define 
parameters such as model order, which can profoundly affect the behaviour 
of linear estimators. Results from the use of different model orders 
should be interpreted with extreme caution. Second, transient interactions 
cannot be defined by this approach. Unstable states of entrainment are 
characterised by significant non-stationarity in the response of a system, 
during which relationships change from cycle to cycle and even 
autoregressive spectral estimation will fail. The frequency tracking locus 
(as described below) is specifically designed for these conditions and can 
give cycle by cycle descriptions of the phase-amplitude parameter space 
and its variation with time. 
The rhythmical neuronal activity responsible for spontaneous respiratory 
drive and the effects upon it of periodic lung inflations have been 
modelled using forced, non-linear equations. Where patients are allowed to 
breathe spontaneously during mechanical ventilation, such as in the sick 
newborn infant, stable entrainment is difficult to achieve. Studies of 
interactions between spontaneous respiratory activity and mechanical 
ventilators in adult humans and in animal models have revealed that 
entrainment of spontaneous respiration by the ventilator stimulus can 
occur under favourable conditions. In adults and animals entrainment 
appears to be induced by the activity of reflexes: the Hering-Breuer 
inflation reflex (which shortens spontaneous inspiration when inflation 
occurs during inspiration) and the Hering-Breuer deflation reflex (which 
lengthens expiration when inflation occurs during spontaneous expiration). 
Vagotomy abolishes entrainment, demonstrating the essential role of 
pulmonary reflexes mediated by parasympathetic pathways. Respiratory 
reflexes similar to those inducing entrainment in adults are present in 
the neonate. 
SUMMARY OF THE INVENTION 
According to a first aspect of the present invention there is provided a 
monitor for monitoring the physiological status of a subject connected to 
an apparatus for providing the subject with artificial physiological 
stimulation, the monitor comprising first transducer means arranged to 
produce a first signal representative of a physiological status of the 
subject, second transducer means arranged to produce a second signal 
representative of a cyclical status of the said apparatus, and processor 
means arranged to receive said first and second signals and to produce an 
output signal dependent upon the relative phase of said first and second 
signals. 
According to a second aspect of the present invention there is provided a 
respiration monitor for monitoring the respiration of a subject on a 
mechanical ventilator, the monitor comprising first transducer means 
arranged to produce a first signal representative of a respiration status 
of the subject, second transducer means arranged to produce a second 
signal representative of a cyclical status of the mechanical ventilator, 
and processor means arranged to receive said first and second signals and 
to produce an output signal dependent upon the relative phase and/or 
amplitude of said first and second signals. 
According to a third aspect of the present invention there is provided a 
method of monitoring the physiological status of a subject connected to an 
apparatus for providing the subject with artificial physiological 
stimulation, the method comprising using first transducer means to produce 
a first signal representative of a physiological status of the subject, 
using second transducer means to produce a second signal representative of 
a cyclical status of the said apparatus, and using processor means to 
produce an output signal which is dependent upon the relative phase of 
said first and second signals. 
According to a fourth aspect of the present invention there is provided a 
method for monitoring the respiration of a subject on a mechanical 
ventilator, the method comprising using first transducer means to produce 
a first signal representative of the respiration status of the subject, 
using second transducer means to produce a second signal representative of 
a cycle status of the mechanical ventilator, and using processor means to 
produce an output signal which is dependent upon the relative phase of 
said first and second signals. 
All of the compatible features set out in the specific description and the 
claims (not directly related to respiration) can be used with the 
invention in its broadest sense. In particular, the frequency-tracking 
locus can be used for many other cyclical physiological signals (eg 
heartbeat etc). 
In one preferred arrangement, a feedback mechanism is provided whereby the 
frequency of the ventilator is controlled in dependence upon the relative 
phase of the first and second signals, and preferably in dependence upon 
the value for the time being of the path length index (PLI). 
If the mechanical ventilator with which the apparatus of the present 
invention is to be used is capable of intermittent mandatory ventilation 
(IMV) the clinician can determine the frequency of mechanical ventilation 
for 1:1 entrainment with the subject's spontaneous respiratory efforts. 
The prediction determined by the spontaneous inter breath inverval (IBI) 
during IMV has been found to substantially more accurate than the IBI 
without any mechanical ventilation at all. Accordingly, in this aspect, 
the invention has improved the ability to predict the rate at which a 
mechanical ventilator should be set in order to achieve a prolonged state 
of 1:1 entrainment with spontaneous respiratory efforts. 
A further advantage is that ventilation is continued at a low rate (using 
IMV) while the necessary calculations are performed, thus avoiding the 
deterioration in condition caused by the previously known method of 
discontinuing mechanical ventilation entirely in order to calculate 
spontaneous breathing frequency. 
Preferably, the onset of spontaneous inspiration is detected, and this is 
used to create a signal which triggers the inflation phase of the 
ventilator. 
The major problems of patient-triggered ventilation are: 
1) inability of patient to "trigger" ventilator due to respiratory muscle 
fatigue, extreme prematurity, reduced central drive to respiration, 
2) "autotriggering" i.e. mechanical inflations initiated by "noise" rather 
than by actual patient effort, and 
3) delay in onset of mechanical inflation following the start of 
spontaneous inspiratory effort, usually due to inadequate sensor 
placement, system delay (physical, biological and electronic). 
The preferred monitoring system can: 
a) assess phase angle between trigger events and the onset of mechanical 
inflation (used to adjust sensitivity and sensor placement/type), 
b) detect "autotriggering" by showing loss of respiratory effort and 
excessive regularity of spontaneous respiration, and 
c) reveal inadequate detection of spontaneous inspiratory effort due to 
inappropriate settings of gain, or very irregular respiratory effort which 
may indicate airway blockage, or the need for sedation.

DESCRIPTION OF PREFERRED EMBODIMENTS 
Referring now to the drawings, FIG. 1a shows in schematic form a comparison 
of the output and stimulus signals occurring in a nonlinear system in 
which the two signals vary in relative amplitude and phase difference on a 
cycle by cycle basis. For each cycle of the stimulus and output, an 
amplitude ratio M.sub.n, and a phase difference .THETA..sub.n are 
obtained. M.sub.n is defined as the ratio of the amplitude of the output 
waveform to the input waveform for each cycle, in other words: 
##EQU1## 
The values of M.sub.n and .THETA..sub.n are plotted as phasors on an 
Argand diagram (FIG. 1b). The first phasor starts at the origin, and each 
subsequent phasor is then plotted from the tip of the preceding phasor. 
Thus, each cycle is represented as a phasor describing the relative 
amplitude and phase of the output to the input signal. Under conditions of 
stable entrainment the individual cycle phasors will align. With a reduced 
influence of the stimulus, entrainment will become unstable and the 
individual phasors will vary both in length and direction. 
In order to quantify the degree of departure from stable entrainment a 
single parameter (the path length index) is used. The path length index 
(PLI) is calculated as follows: 
##EQU2## 
Where M.sub.i =the length of the line from the origin to the tip of the 
terminal phasor. 
For stable entrainment the path length index will be close to unity, and 
the index will progressively increase as the degree of chaos increases. 
In the context of the present invention, the path which is plotted out by 
the phasors on the Argand diagram of FIG. 1b will be known as the 
"frequency tracking locus". A real time plot of the frequency tracking 
locus, for example on a computer screen, can provide the clinician with 
valuable information concerning the particular system being monitored, in 
particular in the present example with information on the interactions 
between the spontaneous respiration of a patient and the respiration 
provided by means of a mechanical ventilator. 
In order to test the invention, the frequency tracking locus was applied to 
data recorded from two preterm infants undergoing mechanical ventilation 
during intensive care. Airway pressure was measured at the proximal 
airway, and oesophageal pressure was measured using a 4 cm long balloon in 
the mid-oesophagus, both being transduced via Gaeltec pressure 
transducers. Signals were band-pass filtered and were recorded on FM tape 
for later digital analysis. Airway pressure (FIG. 2a) represents the input 
to the non-linear oscillatory activity of the respiratory centre, and is 
deterministic, the babies in both examples being ventilated using a 
time-cycled, pressure-limited ventilator at a constant rate and pressure 
which was determined by clinical considerations. Oesophageal pressure 
(FIG. 2b) represents the output from the process, the raw data consisting 
of elements derived from passive ventilation in the absence of spontaneous 
respiratory effort, and spontaneous respiration itself. In order to 
separate the effects of active inspiratory effort from those of passive 
lung inflation an ensemble average of oesophageal pressure data was 
calculated over 100-500 airway pressure cycles. This ensemble was 
subtracted from the raw oesophageal pressure data to give the best 
estimate of active inspiratory effort, as shown in the reconstructed 
oesophageal pressure signal (FIG. 2c). 
Further details of the preferred method used for separating out the effects 
of active inspiratory effort from those of passive lung inflation will be 
described later. It should be noted that all the signals shown in FIGS. 
2a-2c have been low pass filtered. 
Once the signals have been ensemble averaged and low pass filtered, 
calculation of the frequency tracking locus can begin. The airway pressure 
signal is taken as the input signal of FIG. 1a, and the reconstructed 
oesophageal pressure signal as the output signal. The values of M.sub.i 
and .THETA..sub.i are then determined on a cycle by cycle basis. 
Because neither the input signal nor the output signal are simple 
sinusoidal curves, appropriate detection algorithms are required to 
determine the defined points on the input and output signals between which 
the angle .THETA..sub.i is to be measured. One possibility is for the 
breath detection algorithm to select the maximal rate of fall in the 
reconstructed oesophageal pressure signal which is nearly coincident with 
the onset of diaphragmatic contraction (that is, the onset of spontaneous 
inspiration). The timing of this event is compared with the rise in airway 
pressure which marks the onset of inflation by the mechanical ventilator. 
The onset of spontaneous inspiration is compared with the onset of 
mechanical inflation which occurs within one half period of the ventilator 
cycle, and is converted to the phase angle .THETA..sub.i, in degrees, as 
follows: 
##EQU3## 
Where tV.sub.i is the onset of mechanical inflation, tI.sub.i is the onset 
of spontaneous inspiration, and 
##EQU4## 
Determining the maximum rate of fall of the reconstructed oesophageal 
pressure signal is only one method of identifying the commencement of a 
cycle. In the preferred embodiment of the invention, the apparatus allows 
the user to select four separate algorithms to determine the start of each 
cycle. The methods are as follows: 
(1) maximum turning point: the start of each cycle is identified simply as 
the point at which a maximum turning point occurs. 
(2) minimum turning point: the start of each cycle is identified as the 
point at which a minimum turning point occurs. 
(3) Downstroke: the maximum negative rate of change (maximum downstroke) 
over the entire signal sample is determined. The start of each cycle is 
then defined as the point at which the negative rate of change of the 
signal exceeds a user-defined percentage of the maximum downstroke. 
(4) Upstroke: the maximum positive rate of change (maximum upstroke) over 
the entire signal sample is determined. The start of each cycle is then 
defined as the point at which the positive rate of change of the signal 
exceeds a user-defined percentage of the maximum upstroke. 
In all cases, an amplitude rejection threshold is set by the user. This 
specifies a percentage of the mean of all individual cycle amplitudes. Any 
cycle whose amplitude is below this level is subsequently discounted and 
regarded as merely an unwanted minor turning point. In this way, any small 
dips or peaks in the signal may be ignored; for example, the small dips in 
FIG. 2c, between the main dips. 
The reference signal (the airway pressure in this case) normally has a 
steady amplitude that enables the beginning of each cycle of the reference 
signal to be easily identified. A simple way of doing this is merely to 
define a threshold, and the beginning of each cycle is then determined as 
the point at which the signal crosses this threshold in a positive 
direction. The threshold may be set as a user-defined percentage of the 
maximum amplitude (for example, 10% or 20% of the maximum airway pressure 
amplitude). 
Three examples will now be described, showing how the frequency tracking 
locus differs for stable and unstable states of baby-ventilator 
interaction. 
EXAMPLE 1 
Stable Interaction 
During a stable period of mechanical ventilation (FIGS. 3a and 3b) the 
airway pressure signal (FIG. 3a) and the major downward deflections of the 
reconstructed oesophageal pressure record (FIG. 3b) are seen to be in a 
constant phase relationship. The frequency tracking locus over this period 
of 32 seconds reveals that the phasors lie close to the line of ideal 
entrainment, although minor phase changes still occur (FIG. 3c). The 
overall direction of the frequency tracking locus is at +125.degree. (by 
convention, the output signal is said to "lead" the input signal by 125 
degrees). The path length index is 1.034. Each ventilator inflation is 
accompanied by a spontaneous respiratory effort at a fixed phase 
relationship (1:1 entrainment). Noise in the oesophageal signal will be 
detected by the breath detection algorithm when its amplitude exceeds a 
given proportion of a true spontaneous inspiratory effort. Even when noise 
is wrongly interpreted as a spontaneous respiratory effort, however, the 
length of the phasor will be small, and contribute little to the overall 
path length index value. 
EXAMPLE 2 
Chaotic Interaction 
During highly unstable ventilation there is no readily discernible 
relationship between airway pressure and oesophageal pressure deflections 
(FIGS. 4a and 4b), although rapid changes in the rate of spontaneous 
breathing are apparent. The frequency tracking locus demonstrates three 
major features (FIG. 4c) i.e. phase jumps, relatively stable regions (e.g. 
region between points A and B) and chaotic regions. Variations in the 
length of the individual phasors are related to changes in the extent of 
spontaneous respiratory effort. The path length index is 4.172. Changing 
phase relationships reveal a lack of entrainment of spontaneous 
respiratory effort by the ventilator stimulus. In the example shown, there 
is a short period of unstable entrainment of spontaneous respiration by 
the mechanical ventilator between points A and B (arrowheads), which 
occurs in the overall direction of +90.degree. and persists for seven 
spontaneous respiratory cycles. 
EXAMPLE 3 
Integer Ratio Entrainment During Low Rate Mechanical Ventilation 
In this example an examination of the signals (FIGS. 5a and 5b) reveals 
that 32 spontaneous respiratory cycles are associated with 16 ventilator 
airway pressure cycles. Changes in both amplitude and in spontaneous 
respiratory rate can, however, be seen and are clearly seen in the 
frequency tracking locus description of the signals (FIG. 5c). The 
frequency tracking locus reveals an unstable 2:1 interaction where the 
overall direction of the locus varies and small changes in phase angle 
occur with each spontaneous respiratory effort. The path length index is 
4.364. A "zig-zag" course is typical of 2:1 interactions. 
Typically, neonates are ventilated using pressure-limited, time-cycled 
ventilators at a fixed rate. During this form of ventilation, several 
different responses to individual inflations have been documented, namely 
apnoea, augmented inspiration, synchonous breathing, active expiration and 
reflex inhibition of inspiration. The effects of some of these types of 
response to mechanical ventilation upon the frequency tracking locus can 
be predicted. Apnoea or synchronous interaction will lead to an area of 
stability in the frequency tracking locus, which will tend to approach the 
line of ideal entrainment (e.g. FIG. 3c). Consecutive phasor lengths will 
be similar, as spontaneous respiratory amplitudes will be consistently 
low. The path length index will approach unity for these regions. The 
phase angle describing the frequency tracking locus will be close to zero 
degrees for apnoea and passive ventilation. We have found that 1:1 
entrainment only occurs between ventilator stimulus and spontaneous 
respiratory effort in an individual infant during mechanical ventilation 
at a particular rate within a narrow range of phase angle: i.e. the 
overall direction of the frequency tracking locus will be constant during 
1:1 entrainment, but the direction of the frequency tracking locus during 
1:1 entrainment will be specific to an individual baby over a finite 
period e.g. example 2 (shown in FIG. 4c, between points A and B). When 
other phase relationships are apparent, 1:1 entrainment is not maintained. 
Reflex activity (or sudden, unpredictable activity such as gasps and 
hiccoughs) can be represented as a phase jump in the relationship between 
spontaneous respiratory activity and mechanical inflation. The path length 
index will be greater than unity for these areas. Reflex activity which 
affects the duration of spontaneous inspiration or expiration over 
multiple respiratory cycles (such as that occurring in Example 3, FIGS. 
5a-5c) can be responsible for repetitive structures in the frequency 
tracking locus, but these are likely to be short-lived as the relative 
strength of each reflex is dependent upon several factors. Example 3 shows 
an expiratory prolongation reflex which regularly interferes with the 
spontaneous respiratory rhythm to induce a lengthening of the interbreath 
interval associated with each ventilator inflation (FIGS. 5a and 5b). This 
reflex interaction causes a 2:1 interaction. 
In the above illustrations, the path length index is little affected by 
changes in relative magnitude when phase relationships are unvarying from 
cycle to cycle. The major effect on the path length index comes from 
changes in relative phase angle, and it is thus ideally suited to 
providing information about entrainment phenomena when stimulus and 
resultant frequencies are similar. When frequencies are widely different, 
the frequency tracking locus can describe polygonal paths and, thus, a 
very high path length index value can be obtained. 
Phase jumps reflect the reflex response of one waveform (spontaneous 
respiratory effort) to individual cycles of the stimulus (the airway 
pressure signal), and the frequency tracking locus permits the recognition 
of reflex phenomena forming phase discontinuities. The frequency tracking 
locus is thus ideally suited to the description of respiratory entrainment 
phenomena in the clinical setting where the response to individual 
ventilator inflations and overall ventilatory strategy has to be assessed. 
The system described herein gives clinicians a measure of the degree to 
which spontaneous breathing efforts by the patient are matched by the 
actions of the mechanical ventilator from breath to breath. The complex 
interactions between the patient and the ventilator are broken down to a 
single number representing the degree of chaos in the relationship between 
the patient's own breathing and the action of the mechanical ventilator. 
In the preferred embodiment, this number, the Path Length Index (PHI) is 
recalculated every 16 to 64 seconds, and is then displayed as a trend over 
several hours. The trend of the PLI can then be used to assess the effects 
of treatment (such as sedative drugs used to suppress the patient's own 
breathing) and changes in ventilation strategy (for example 
patient-triggered ventilation) upon the pattern of ventilation achieved. 
The system may also display, for example on a computer screen, the progress 
of the frequency tracking locus (see FIG. 1b) over a period of several 
hours. The airway pressure and the reconstructed oesophageal pressure may 
also be displayed as individual traces, with or without low band-pass 
filtering and/or removal of noise and base line effects. 
Turning now to FIG. 6, there is shown in schematic form an embodiment of 
apparatus in accordance with the present invention. A first transducer T1, 
which may be attached to a subject, produces a first electrical output 
signal dependent upon the respiration status of the subject. A ventilator 
V includes transducer means T2 to produce a second electrical signal which 
is dependent upon the ventilator inflation. The transducers are coupled to 
a processor P which includes signal comparator means to produce first and 
second control signals dependent upon the relative phase of signals from 
the first and second transducers and the ratio of the amplitudes of these 
signals. The display D is fed with the control signals and produces, 
amongst other things, a display of the frequency tracking locus of the 
transducer signals. As indicated above, the display may also show the raw 
transducer signals and/or the signals after filtering and removal of noise 
and base line effects. The incoming signals, specifically from the 
transducer means T1, will normally be low pass filtered to remove high 
frequency noise; that may be carried out by the processor P. 
Optionally, the output of the processor P may be coupled to the ventilator, 
as shown by the dotted line in FIG. 6, so as to produce a feed-back path. 
In this way, the operation of the ventilator can be controlled in 
dependence upon the path length index, and perhaps also in dependence upon 
other features of the signals produced by the transducer means T1 and T2. 
We have described a new form of measurement of total, chaotic activity in a 
system in which the stimulus and resultant signals are described in terms 
of phase and relative amplitude, cycle-by-cycle. We have demonstrated the 
application of this technique to a clinical situation in which complex 
interactions are known to be a prominent feature. The Frequency Tracking 
Locus can be used to provide a qualitative description of baby-ventilator 
interactions, and a quantitative description of changes in relative phase 
and amplitude (the Path Length Index) which corresponds to an "Index of 
Favourability" of ventilator strategy. This method allows clinicians 
caring for newborn infants to optimise mechanical ventilation for the 
individual by estimating the response to changes in therapy, such as the 
use of different ventilator rates, sedative agents and patient-triggered 
ventilation. 
In the embodiment so far described, the output signal has been the 
reconstructed oesophageal pressure of the patient. Other methods for 
detecting spontaneous respiration could also be used, including inductance 
plethysmography, transthoracic impedance pneumography, diaphragmatic 
electromyogram and flow/volume signals from pneumotachography. Another 
convenient method is to attach an abdominal movement sensor (such as that 
used in apnoea monitoring). In each case, the output signal is then 
compared with the airway pressure signal from the ventilator circuit to 
the patient end, which is already available as an output from many modern 
ventilators. 
In the embodiment of the apparatus in which there is no feedback from the 
processor to the ventilator (in other words where the dotted line shown in 
FIG. 6 is absent) it is necessary for the clinician to decide on the 
operating frequency of the ventilator. The optimal form of baby-ventilator 
interaction is generally thought to be a stable 1:1 relationship between 
spontaneous respirations and ventilator inflations so that each of the 
babies' breaths is accompanied by an artificial lung inflation. The 
difficulty, in the past, has been determining exactly which frequency to 
choose to produce the desired 1:1 entrainment. Previous studies have used 
a short period of disconnection from the mechanical ventilator in order to 
determine the spontaneous, unstimulated rate of respiration by the baby. 
This method does not take account of the interactions between the baby's 
own respiratory effort and the train of mechanical inflations. These 
interactions result from reflex activity caused by the stimulation of 
receptors within the lungs and chest wall when inflation occurs. 
Interactions are made more complex by the introduction of "noise" into the 
system from spontaneous motor and respiratory activity associated with 
clinical procedures and pain, and from unpredictable activity such as 
crying, gasps, sighs and hiccoughs. 
We have found from our research that it is possible to calculate the 
frequency at which 1:1 entrainment will take place by measuring the 
interval between spontaneous breaths during intermittent mandatory 
ventilation (IMV). Intermittent mandatory ventilation is illustrated in 
FIGS. 7a and 7b, the upper trace of which shows the airway pressure as 
forced by the ventilator at regular intervals. The lower trace shows the 
oesophageal pressure, and it will be seen that the baby is breathing 
relatively irregularly, but several times during each inflation cycle of 
the mechanical ventilator. 
Once the oesophageal pressure signal has been averaged, the mean frequency 
is found to be a good guide as to the ventilator frequency at which the 
desired 1:1 entrainment will occur. Typically, the average frequency of 
spontaneous breath during IMV is not the same as the average frequency in 
the absence of external ventilation. 
Any of the other alternative ways of determining active breathing, 
mentioned above, could also be used. 
We will now describe the results of a research study which confirm that, 
with the present invention, the system has the ability to predict the rate 
at which a mechanical ventilator should be set in order to achieve a 
prolonged state of 1:1 entrainment. 
20 infants were studied on 35 occasions during intermittent mandatory 
ventilation (IMV) when they were recovering from idiopathic respiratory 
distress syndrome. Orotrachael tubes delivered regular inflations at rates 
of .ltoreq.34/min via pressure-limited, time-cycled neonatal ventilators 
with inspiratory times and rates determined by clinical staff on the basis 
of blood gas measurements. The study was approved by the District Ethics 
Committee. Written parental consent was obtained in all cases. 
Spontaneous respiratory timings were measured from an abdominal pressure 
capsule and, in some cases, compared with timings derived from oesophageal 
pressure measurements, transthoracic impedance, and surface diaphragmatic 
EMG. Oesophageal pressure was measured in the mid-oesophagus according to 
the method of Beardsmore et al. Oesophageal pressure measurements were not 
analysed when excessive peristaltic activity obscured the effects of 
spontaneous respiration. Airway pressure was measured from the proximal 
airway and abdominal capsule measurements were performed by using the 
Wright apnoea monitor according to the description of South. Diaphragmatic 
EMG signals were recorded on infants from bilateral skin electrodes placed 
over the anterior costal margin in right and left midclavicular lines. ECG 
artefact was subtracted using an R-wave triggered gate of length 40-100 
msec and the output was via an root mean square (RMS) integrator with a 
time constant of 200 msec. Feedback from the RMS integrator output to the 
logic gate meant that the gate length window in the data was filled by 
averaged RMS output. The raw signal was delayed by 20-50 msec to centre 
the ECG complex in the gate period. The resultant diaphragmatic signal is 
referred to as RMS-EMG. Pulse oximeter values of arterial oxygen 
saturation and plethysmographic waveform were monitored throughout 
recordings (Nellcor N200). 
Recordings were made during periods when infants were not crying and when 
gross motor movements were minimal. All infants were sedated with an 
opiate (pethidine) or chloral hydrate, given regularly as part of routine 
care within six hours of the start of recording. No attempt was made to 
formalise sleep staging because of the low gestational age of the infants 
and the presence of sedative agents which made meaningful descriptions of 
sleep state impossible. A period of 15 minutes was allowed between 
handling and the start of recording. Data were recorded on magnetic tape 
for periods of up to 2 hours. Analysis was performed off-line when 
recorded signals were played back through a low-pass filter (6 Hz cut-off) 
and underwent analogue-digital conversion of 16 Hz sampling rate. Arterial 
blood gases were taken shortly before the start of each recording and 
arterial PaCO2 was always between 35 and 45 mmHg. Arterial oxygen 
saturations lay between 90 and 96% for duration of recordings. No babies 
exhibited an arterial PH&lt;7.35. 
The most important methodological aspect of the study was the use of random 
fluctuations in spontaneous respiratory rate occurring against a 
background of deterministrically applied lung inflations at constant rate 
as shown in FIGS. 7a and 7b. Vertical bars in FIGS. 7a and 7b mark the 
inspiratory-expiratory juncture associated with each of the first four 
ventilator inflations: inflation occurs during both spontaneous 
inspiration and expiration. It will be seen, therefore, that lung 
inflation occurs at all phases of the spontaneous respiratory cycle so 
that the reflex response of the infant's respiratory system is 
interrogated during inspiration, expiration and also during transitions 
between these phases. 
In order to separate the passive effects of mechanical lung inflation upon 
AC and OP (caused by the effects of inflation of the lungs causing a 
transmitted positive pressure) from the effects of spontaneous respiratory 
effort (caused by changes induced by descent of the diaphragm), an 
ensemble average of each of the waveforms was built up over 100-1000 
ventilator inflations. FIG. 8 shows a typical ensemble average over 300 
ventilator cycles of airway pressure (bold line) and abdominal capsule 
(dotted line). On average, there is a passive effect of inflation upon the 
abdominal capsule signal which then returns to the base line at a rate 
determined by the effective time constant of the circuitry. Further, 
damped oscillations occur which are due to interference by the cardiac 
impulse. As may be seen from FIG. 8, the abdominal capsule trace is nearly 
coincident with the airway pressure trace during mechanical inflation. The 
random distribution of spontaneous respiratory effects relative to 
ventilator inflation has led to the averaging out of spontaneous effort so 
that only the regular, passive component of the abdominal capsule signal 
due to mechanical inflation has remained. A similar result (not shown) is 
obtained if one determines the ensemble average of the oesophageal 
pressure trace. After calculation of the ensemble average the average was 
subtracted from the raw data for each ventilator inflation period in order 
to give an estimate of "active" respiratory components only (FIGS. 2a-2c). 
In the subsequent description, the oesophageal pressure trace will be known 
as the OP trace, the abdominal capsule trace as the AC trace and the 
airway pressure trace as the AP trace. 
In order to ensure that the ensemble averaging subtraction did not 
materially affect spontaneous respiratory timings in OP and AC traces when 
spontaneous breaths occurred in association with mechanical inflation, we 
compared the RMS-EMG signal with those of OP and AC before, and after 
subtraction of the ensemble average. Although mechanical (AC and OP and 
"neural" (RMS-EMG) respiratory timings were not identical, reflex 
prolongation of expiration was apparent in the RMS-EMG signal as well as 
in the reconstituted OP and AC signals: this phenomenon was not an 
artefact induced by the digital averaging and subtraction processes. 
After subtraction of the passive, inflation-related component from OP and 
AC signals, AP OP and AC data were subjected to microcomputer-based 
algorithms which determined the onset and duration of mechanical 
inflation, the onset and end of each spontaneous inspiratory effort, and 
the intervals between all these events. 
Spontaneous inspiratory duration, Ti, was defined as the period between the 
onset of a sharp fall in OP (or rise in AC) signal and the next minimum 
(maximum) in the signal provided that the apparent amplitude of 
inspiratory effort was greater than a preset limit. This limit was 
determined over the entire recording by automated algorithms. Spontaneous 
expiratory duration, Te, was defined as the period between the minimum 
(maximum) in the OP (AC) waveform and the next inspiratory downstroke 
(upstroke). 
The breath detection algorithm selects the maximal rate of fall (rise) in 
the reconstructed oesophageal pressure (abdominal capsule) signal. We have 
confirmed that this event is nearly coincident with the onset of 
diaphragmatic EMG activity from careful examination of raw signals 
digitised at a sampling rate of 2000 Hz. The timing of this event was 
compared with the rise in airway pressure which marked the onset of 
inflation by the mechanical ventilator. In this way it was possible to 
differentiate between those spontaneous respiratory cycles any phase of 
which were associated with a ventilator inflation, from those spontaneous 
respiratory cycles which started and ended in the absence of any 
ventilator inflation. 
All channels of reconstituted data (after subtraction of ensemble average) 
were processed in segments of 256 samples, which is the equivalent of 16 
seconds of data at 16 Hz sampling rate. 
Two methods were used concurrently in order to achieve maximum accuracy in 
the recognition of true, spontaneous respiratory effort. 
(1) The first difference of the sampled data series, .delta.y, was 
interrogated for each segment of 256 samples in order to derive the mean 
.delta.y for that segment. The segment was then scanned For regions of 
.delta.y in excess of a user-defined percentage (typically 10-20%) of mean 
.delta.y for the segment. 
(2) The threshold for distinction between noise and a true spontaneous 
respiratory effort in the signal was set by detection of the turning 
points (maxima and minima) and calculating the difference in magnitude 
between successive turning points for the whole data segment. 
For AC,TTI, and inductance plethysmography, a breath was deemed to have 
occurred when the first difference, .delta.y, of the sampled data series 
was greater than a predefined percentage of the mean for that data segment 
and when the difference between the onset of inflation and the next 
maximum encountered is greater than a predefined proportion (typically 
0.5-0.6) of the mean difference between successive maxima and minima as 
defined in (2) above. 
For oesophageal pressure signals, where spontaneous inspiration is marked 
by a sudden fall in signal amplitude, a breath is deemed to have occurred 
when the first difference, .delta.y, of the sampled data series is less 
than a predefined percentage of the mean for that data segment and when 
the difference in magnitude between the onset of inspiration and the next 
minimum encountered is greater than a predefined proportion (typically 
0.5-0.6) of the mean difference between successive maxima and minima as 
defined in (2) above. 
For the RMS-EMG and airway pressure signals, detection of spontaneous 
inspiration or mechanical inflation occurs when the signal attains an 
arbitrary, predefined value above the noisy baseline. The end of 
spontaneous inspiration or mechanical inflation is defined as that point 
when the signal falls below this threshold value. The maximum occurring in 
this region of the signal was then compared with the threshold for 
acceptance as a true breath or inflation according to (2) above. 
Expiration is defined in all spontaneous respiratory channels as being the 
period between the end of one spontaneous inspiration and the onset of the 
next i.e. as the non-inspiratory period. 
The major sources of artefact in the recognition of spontaneous inspiratory 
activity in the different data channels are: 
(a) electromechanical noise (all channels) which causes baseline shifts and 
random fluctuations in respiratory timings: artefacts were avoided by 
averaging processes and setting thresholds for recognition of real events 
so as to exclude baseline shifts; 
(b) spontaneous motor activity (all channels except AP) can be mistaken for 
respiratory effort: not specifically excluded, but avoided by setting 
upper limits of plausibility for respiratory timings; 
(c) peristaltic activity and changes in physical disposition of sensor (OP 
only)--inherent problem in oesophageal ressure traces: avoidance as in (b) 
above; 
(d) very rapid rates of respiration (all channels): rates above 3 Hz (rare) 
will be beyond the resolution of this system due to sensor and signal 
processing constraints; 
(e) damped oscillations following activity (all channels)--due to filtering 
and "AC coupling" circuitry: avoided by use of ensemble averaging 
technique which displays, calculates and subtracts those effects from raw 
data signals (for all save DEMG); 
(f) effects of analogue and digital filtering (all channels): as (d) above; 
(g) sampling error (all channels): avoided by low-pass filtering at cut-off 
below Nyquist frequency; and 
(h) excessive humidity (AP only): clinical recognition. 
Student's paired t-test was used to compare mean values of Ti, Te and IBI 
which were associated/unassociated with the inflation time of the 
mechanical ventilator for the duration of each recording in each subject. 
Results--Effects of Lung Inflation on Respiratory Timings 
Baseline Te was 0.48 s (.+-.0.129) and increased to 0.65 s (.+-.0.182, 
p&lt;0.0001) when inflation was coincident with spontaneous respiration. 
Baseline Ti was 0.34 s (.+-.0.062 and increased to 0.38 s (.+-.0.081, 
p&lt;0.0001) with inflation. IBI increased from 0.82 s (.+-.0.161, equivalent 
to 73/min) to 1.03 s (.+-.0.201, equivalent to 58/min) with inflation, an 
increase of 26%. 
Prediction of Frequency for 1:1 Entrainment 
On several occasions, an increase in the rate of mechanical inflation to 
50-60/minute resulted in 1:1 entrainment, characterised by a fixed phase 
relationship between inflation and spontaneous respiration in which 
mechanical inflation occurred at the end of inspiration, or early in 
expiration i.e. inflation and inspiration were out of phase (FIGS. 3a-3c). 
Unpredictable activity, such as gasps, were rapidly followed by a return 
to the phase-locked pattern. 1:1 entrainment of this type was seen for 
several hours. As ventilator rates fell with improved lung function, 
studies during IMV revealed that the rate of inflation to induce "stable" 
entrainment could only be predicted from the inter-breath interval (IBI 
(equivalent, on average, to a rate of 58/min) associated with inflations: 
during continuous positive airways pressure (CPAP) in the absence of 
mechanical lung inflations the spontaneous rate would be considerably 
higher (equivalent to 73/min, the overall average IBI of unassociated 
breaths above). 
1:1 entrainment of spontaneous respiration could be induced at a rate which 
was predicted from the IBI during IMV, but not from the spontaneous 
respiratory rate in the absence of mechanical lung inflations. An example 
of an infant in a 1:1 entrained state at 57 inflations per minute who was 
switched to continuous positive airways pressure (CPAP) without inflations 
is given in FIGS. 9a and 9b. Note the immediate reversion to a higher, 
unstimulated rate of breathing of 79 breaths per minute when inflations 
are halted. 
Pulmonary reflexes in the newborn have usually been investigated by the 
occlusion technique, and results from these studies may not be comparable 
with those from studies of lung inflation because of the effects of chest 
wall distortion on the intercostal phrenic inhibitory reflex. Despite 
early reports that the Hering-Breuer expiratory prolongation reflex 
diminishes with postnatal age, and was less active in preterm infants, 
more recent work confirms the findings of our study i.e. that expiratory 
prolongation reflexes are present even down to the lowest gestational ages 
and do not diminish over the first days and weeks of life. 
Our results show that cases IBI was prolonged in all cases by the effects 
of mechanical inflation. IBI changes were dominated by changes in Te, and 
a strong correlation existed between IBI and Te in most of the cases 
studied. IBI was more prolonged for inflations occurring during the 
spontaneous expiratory phase than during spontaneous inspiration, the 
extent of IBI prolongation being proportional to the duration of 
expiration at the instant when mechanical lung inflation occurred. As 
spontaneous respiratory rate was largely determined by changes in Te, 
greater prolongation of IBI occurred when spontaneous respiratory rate was 
lowest. 
Changes in Ti were less marked, but Ti usually increased when inflation 
occurred during the spontaneous inspiratory phase. This finding is in line 
with other studies of spontaneous respiratory timing in preterm infants, 
where the response to lung inflation does not switch off inspiratory 
drive, as it would in the adult, being apparently mediated in adults by a 
slowly reacting stretch receptor. The inspiratory prolongation which is 
seen in the preterm infant is thought to be due to the prevention of 
rib-cage distortion by diaphragmatic activity, hence eliminating another 
reflex, possibly the intercostal phrenic inhibitory reflex described by 
Byran et al, the effect of which will be to shorten inspiration. 
Entrainment (phase-locking) of spontaneous respiration by rates of 
inflation less than the spontaneous rate must occur by slowing of the 
spontaneous rate of breathing. Previous descriptions of attempts to induce 
"phase synchrony" between mechanical ventilation and spontaneous 
respiratory activity have usually relied upon the observation of 
respiratory rates when the baby is switched from regular ventilation to 
CPAP for short periods. This method will not allow the effects of a train 
of inflations upon spontaneous respiration to be evaluated. Thus, the use 
of IMV induced reflex activity predicted the rate at which stable 
entrainment of spontaneous respiration could be achieved. Even during 
"stable entrainment", minor changes in spontaneous respiratory rate were 
occurring, but the effects of inflation-induced IBI prolongation reflexes 
were sufficient substantially to correct the subsequent IBI and maintain a 
1:1 relationship. 
It is clear from the above discussion that entrainment of spontaneous 
respiration in this way is only possible by rates of mechanical 
ventilation which are less than the spontaneous rate when unaffected by 
inflations. 
As will be clear from the preceding descriptions, the apparatus embodying 
the present invention will desirably be capable of producing intermittent 
mandatory ventilation (IMV) for a certain period, and during that time 
have means for calculating the spontaneous inter-breath interval (IBI) of 
the patient. The apparatus should then either be automatically or manually 
switchable to produce continuous ventilation at a frequency at or about 
the IBI which has previously been determined. During an initialisation 
phase (for example ten minutes) the apparatus determines the average 
relationship between the input and output signals, following which the 
frequency tracking locus is displayed, along with the path length index 
(PLI) and the trend of the PLI. The input and output signals may also be 
displayed, after the subtraction of base line effects and noise. The 
averaging process may automatically be repeated, at defined intervals, for 
example every ten minutes.