Respiratory rectification

ES 2 559 933 T3

DESCRIPTION

Respiratory rectification.

Background

[0001] In humans, gas exchange is accomplished by rhythmic inflation and deflation of the lungs.

During respiratory movements, the lung is passive and its volume is determined by the size of the thorax, which in turn depends mainly on the activity of the intercostal muscles and the diaphragm.

[0002] The vertical movement of the diaphragm is approximately 1.5 cm during quiet breathing and can reach up to 7 cm in deep breathing.

A 1 cm lowering of the diaphragm will increase the thoracic capacity by approximately 300 mL and cause a corresponding amount of air to enter the lungs.

Thus, the movements of the diaphragm can be responsible for approximately 60-80% of the dilation of the lung and the total air breathed.

[0003] A respiratory control center in the brain controls the respiratory muscles.

The respiratory muscles consist of the muscles of the respiratory pump (diaphragm and intercostal muscles) and of the airways.

Neural signals that travel to the respiratory muscles constitute the central respiratory drive. While the diaphragm is the main source of power for the respiratory pump, the function of the airway muscles is to keep the airway open.

Without a central neuronal drive, the airway can collapse or partially occlude.

[0004] During sleep, respiratory control is unconscious and is determined by metabolic demand (mainly the need to remove CO2 from the blood).

In all humans, the central neural drive to the respiratory pump and airway muscles during sleep is reduced compared to a conscious state.

In the presence of a moderate pathology, the resistance of the airways to the flow of air may increase during sleep, resulting, for example, in snoring.

In extreme cases, the airways can close completely resulting in obstructive sleep apnea (OSA). In some cases, a dysregulation of central control can result in periodic breaths and serious illness that can damage multiple organs.

[0005] Central sleep apnea (ACS) is a form of periodic breathing that is characterized by an oscillating central respiratory drive.

ACS can be characterized by a typical increasing and decreasing respiratory pattern composed of alternating apneas and hyperpneas (periods of hyperventilation), historically called Cheyne Stokes respiration (CSR).

[0006] Obstructive sleep apnea (OSA) is characterized by instability of the upper airways.

A collapsed airway prevents or reduces inspiration in the face of increased or continued respiratory effort. A common pattern of OSA in the general population is characterized by periodic stimulation that results in a sudden opening of the airways.

[0007] The inventors have discovered that in some patient populations, such as those with congestive heart failure (CHF) for example, it is difficult if not impossible to separate the underlying mechanisms of OSA and ACS. A purely core CSR Model is rare.

The common presentation of periodic breathing in patients with CHF may consist of alternating respiratory events that may include hyperpneas, hypopneas, and central, obstructive, and mixed apneas.

There is significant overlap, and most patients experience varying degrees of both central and obstructive events.

In fact, it is believed that the oscillating central respiratory drive can lead to closure or increased resistance of the upper airways.

[0008] After an extensive series of experiments, the inventors realized that effective treatment of many cases of periodic breathing required attention to the central neuronal drive to breathe, as well as to airway problems.

Due to the relationship between the respiratory pump muscles and the airway muscles, the inventors wished to modulate the intrinsic respiratory activity of the brain to affect both aspects of respiration.

The respiratory control center of the brain is located in the medulla of the brain and is not directly accessible to modern neuromodulation technologies.

ES 2 559 933 T3

The inventors were therefore forced to investigate neural inputs to the brain that control the behavior of the respiratory control center.

[0009] The respiratory control center of the brain receives inputs from chemoreceptors from the arterial vasculature (in the aortic arch, carotid bodies and blood vessels of the brain itself) and from mechanoreceptors such as the respiratory pump muscles, lung stretch receptors lung, and intercostal stretch receptors.

It is known that altering the entry into the brain of chemoreceptors by having the patient inhale some carbon dioxide can control periodic respiration.

[0010] The phrenic nerves control the movement of the diaphragm, which in turn can lead to the activation of series of neuronal inputs to the brain, for example by stretching various innervated tissues of the thorax.

The Hering-Breuer reflex is a powerful neural feedback from lung stretch receptors to the brain.

When the lung is inflated and stretched, the respiratory center of the brain suspends the breath drive.

Pacing of respiration with external stimuli in sleeping humans is believed to involve the Hering-Breuer reflex as well as other neural inputs.

[0011] The technology that the inventors chose for the research was phrenic nerve stimulation.

A form of phrenic nerve stimulation known as electrophrenic ventilation or diaphragmatic stimulation has been used for many years to replace intrinsic respiration.

The physiology of respiratory stimulation is straightforward.

The phrenic nerve is stimulated to take control of breathing by slightly hyperventilating the patient.

The consequent reduction in the concentration of carbon dioxide in the blood is detected by the respiratory control center of the brain and substantially all the neuronal emission from the center (central respiratory impulse) is stopped.

A patient, thus stimulated, may not experience respiratory disturbances while stimulated but becomes temporarily dependent on phrenic stimulation for ventilation and gas exchange.

[0012] Such phrenic stimulation has been used to treat Paraplegia or Congenital Central Hypoventilation Syndrome in children since the 1960s and successfully prevented death by replacing natural respiration with artificial respiration.

By taking over respiration, such stimulation suppresses the intrinsic central respiratory drive.

In addition to eliminating oscillations in the respiratory drive, they can also eliminate the urge to keep the airway open.

In patients with Congenital Central Hypoventilation Syndrome, this limitation can be overcome through a tracheostomy (a surgical procedure performed on the neck to open a direct airway through an incision in the trachea).

A tracheostomy is unacceptable in patients with periodic breathing.

In contrast, embodiments according to the present disclosure apply stimulation to a single phrenic nerve at a rate that in some embodiments is slightly below the intrinsic rate of the patient.

These embodiments may involve rhythmic pacing of the patient's central respiratory activity rather than intrinsic drive suppression.

When applied during periodic breathing to sleeping individuals and provided within a particular range of stimulation parameters, this form of phrenic nerve stimulation frequently resulted in the restoration of normal breathing rhythm, was sustainable, and was well tolerated during natural sleep.

[0014] In some embodiments according to the present disclosure, the stimulation produces paced contractions of a hemidiaphragm innervated by the stimulated nerve and consequent paced lung inflates.

The patient's intrinsic breathing rate is modulated by stimulation, and it becomes more regular and exhibits less periodicity.

Improved respiration is believed to be achieved through the lung distension mechanism and rhythmic pacing of neurons in the brain's respiratory center, among other mechanisms, rather than through the control of respiration itself.

Spontaneous respiration and central respiratory drive are preferably preserved, thus resulting in the benefit of maintained airway muscle tone and natural regulation of blood gases.

[0015] Another embodiment according to the present disclosure deals with monitoring and control of diaphragmatic stimulation and rhythmic pacing of stimulation.

This embodiment is based on the assumption that when the stimulation rhythmically paces the breath, the spectral power of the respiratory signal in the immediate range of the pulse frequency will be significantly higher than in other frequency bands.

ES 2 559 933 T3

In one embodiment, a ratio of total variance of respiratory signal that falls in the narrow band centered in the pacing rate to the total variance in the wider respiratory rate band can be calculated.

It can be expected that the value of this ratio will increase in proportion to the rhythmic pacing of the stimulation breath, thus allowing a guided and calibrated treatment based on the effect of the stimulation on the actual breath.

Two exemplary calculation methods are described here for this ratio: the spectrum method and the histogram method.

It should be understood that these methods are not the only ways to determine the efficiency of capturing rhythmic or muscle or nerve pacing by frequency analysis, and other methods will occur to those skilled in the art upon reading the present disclosure.

[0016] The inventors observe an occasional regularization of the respiratory rhythm and a resolution of both central and obstructive apnea during unilateral phrenic stimulation.

The inventors observed that phrenic stimulation prevents and corrects the intrinsic periodicity of respiration in patients who frequently show oscillatory breathing during sleep.

The inventors note that the patient's breathing, which was previously highly irregular, gradually paces the rhythm of the phrenic stimulation and follows it if the stimulation index is changed.

The inventors have also observed a fixed and repetitive association between external stimulus and neuronal inspiratory activity in the respiratory control center of the brain.

The inventors confirmed that the patient's breathing was rhythmically paced, and was in fact spontaneous, by making the following observations:

(1) When the stimulation stopped abruptly, the patient did not stop breathing.

In some cases, patients continued to breathe as if they were still being stimulated, following signals that were no longer present for several minutes before returning to the pre-stimulation pattern.

(2) As only one phrenic nerve was stimulated, the second lung was free to inflate and deflate without the direct influence of the stimulation.

The inventors confirmed the independent movement of the unstimulated lung by measurements of independent transthoracic impedance.

(3) Often, patients were observed to insert or intersperse small spontaneous breaths between breaths paced with stimulation, thus confirming that the central respiratory drive remained active.

These interspersed breaths did not interfere with the rhythmic pacing pattern.

[0017] Stabilization of blood gases and elimination or reduction of hypoxia and intermittent stimulations associated with hyperventilation by pacing with phrenic stimulation is believed to improve airway tone.

Hypoxia is an exceptionally strong stimulus to hyperventilate, as is the neurological stimulation that frequently follows hypoxia.

The severity of hyperventilation that follows intermittent hypoxia determines the subsequent reduction of CO2 in the blood that causes withdrawal of neuronal stimulus for both the respiratory pump muscles and the airway muscles.

The opposite also is true.

The reduction of hypoxia and the consequent hyperventilation exacerbated by hypoxia should help maintain the respiratory drive and maintain airway neuronal muscle tone after the end of transient hyperpnea.

[0018] Another way that ACS is believed to reduce upper airway stability and induce OSA is by promoting periods of hypopnea.

During hypopnea, the activity of both the respiratory pump muscles (eg, the diaphragm) and the upper airway dilator muscles (eg, the genioglossus) is reduced.

Therefore, an increasing and decreasing respiratory pattern of the central respiratory drive in an individual with an upper airway prone to collapse may result in obstructive apnea / hypopnea during periods of hypopnea due to upper airway hypotonia ( low muscle tone).

It is reasonable to assume that the opposite is also true.

Reducing hypopnea in ACS patients will help stabilize the airways by increasing airway muscle tone.

[0019] In the context of this disclosure, hypopnea refers generally to a transient reduction in air flow (during sleep) lasting at least 10 seconds due to transient shallow breathing, or an abnormally low respiratory rate.

In the medical literature, breathing that is too shallow (hypopnea) or too slow (bradypnea) is sometimes distinguished.

Hypopnea is less severe than apnea (which is a more complete loss of airflow) but can also

ES 2 559 933 T3 assumes a decrease in the amount of air movement into the lungs and can cause oxygen levels in the blood to drop (hypoventilation).

[0020] There is no firm scientific consensus as to the closed or quantitative definition of hypopnea and it is understood that many definitions are often used in the scientific literature and can be applied.

For example, in their research, which served as the basis for this disclosure, the inventors used a transient reduction in respiration (air flow) of> 30% for a duration of 10 to 60 seconds accompanied by oxygen desaturation. detectable (4%) as a quantitative technical definition of hypopnea.

Within the narrow scope of this quantitative definition, hypopnea and transient hypoventilation are equivalent for practical purposes.

[0022] When broadly defined, hypoventilation is the state in which a reduced amount of air enters the alveoli of the lungs, resulting in decreased oxygen levels and / or increased carbon dioxide levels. in the blood.

Hypoventilation can also be defined broadly, and perhaps better, as breathing that is not adequate to meet the body's needs.

Hypoventilation may be due to hypopnea or reduced lung function.

Hypoventilation can be transient (as a result of hypopnea) or prolonged due to various conditions such as congenital diseases, chronic obstructive pulmonary disease (COPD), or obesity.

[0023] Periodic lung inflates play an important role in maintaining sympathetic-parasympathetic neuronal balance, heart rate, and blood pressure regulation.

As early as the early 1940s many of these physiological interactions were related to neuronal feedbacks signaling lung expansion to the autonomic nervous system.

The role of the autonomic nervous system in the homeostasis of the body is particularly important during sleep.

These beneficial feedbacks have been rated based on lung inflation.

[0024] The importance of neuronal feedbacks from the stretch receptors of the lungs to the multiple centers of the brain that control cardiovascular activity can best be demonstrated by the so-called respiratory sinus arrhythmia.

Respiratory sinus arrhythmia (ASR) is a change in heart rate that occurs in sync with respiration, whereby the RR interval on an electrocardiogram shortens during inspiration and lengthens during expiration.

Although ASR has been used as an index of cardiac vagal function, it is also a physiological phenomenon that reflects universally observed respiratory-circulatory interactions in vertebrates.

Some studies have shown that the efficiency of lung gas exchange is enhanced by ASR, suggesting that ASR may play an active physiological role.

Synchronizing the frequency of alveolar ventilation and its perfusion with the ASR in each respiratory cycle could reduce energy expenditure by suppressing unnecessary beats during expiration and ineffective ventilation during perfusion reflux.

ASR or heart rate variability in synchrony with respiration is a biological phenomenon, which can have a positive influence on gas exchange at the lung level through efficient ventilation / perfusion timing.

[0025] The inventors observed an increase in ASR when patients with CHF were treated according to embodiments of the invention.

It is reasonable to expect that other benefits such as dilation of blood vessels and reduction of malignant arrhythmias will follow.

[0026] Phrenic or diaphragm stimulation in accordance with embodiments of the present disclosure can expand respiration or lung volume to combat the effects of hypopnea.

Stimulation is also expected to result in prolonged lung expansions unlike the rising and falling pattern found in ACS.

An additional benefit for airway tone can also be expected from these improvements.

In patients with significant ACS, upper airway collapse and resulting OSA may be secondary to withdrawal of neuronal and mechanical stimulus to the airways caused by reduced parasympathetic activation and decreased lung inflation during episodes. central apnea / hypopnea.

[0027] The clinical benefit may derive from pacing the intrinsic respiratory rhythm by stimulation.

Increased lung volume, stabilized blood gas composition, and reduced hypoventilation suggest improved airway dynamics in addition to primary correction of central respiratory instability.

ES 2 559 933 T3

The upper airway dilator muscles play an important role in maintaining airway patency.

Many of the pharyngeal dilator muscles are known to demonstrate inspiratory phasic activity, the onset of which precedes diaphragmatic activity.

That is, the airway muscles contract in sync with respiration slightly earlier than the respiratory pump muscles, thus preparing the pharyngeal airway for the development of negative pressure during inspiration.

[0029] The best studied pharyngeal muscle is the genioglossus.

The genioglossus receives information from the respiratory control center of the brain (or more precisely from the central respiratory pattern generator of the brainstem) located in the spinal cord.

The hypoglossal nerve activates the genioglossal nerve, and the hypoglossal nerve has been found to send information 50-100 ms before the phrenic nerve in healthy patients.

[0030] Chemoreceptor inputs are also important in influencing hypoglossal motor nerve emissions.

A low CO2 in the blood (hypocapnia) reduces the activation and a high CO2 (hypercapnia) increases it.

Thus, avoiding hypocapnia and the reduction of airway muscle activation that accompanies hypocapnia may be helpful for any periodic breathing treatment.

[0031] Embodiments according to the present disclosure are usable to treat periodic breathing in patients who are sleeping, but can also be used to regulate breathing in people who are at rest and suffering from ischemic heart disease, heart failure, hypertension, COPD, and other conditions for which improved efficiency of respiration is advantageous.

[0032] In an embodiment according to the present disclosure, a measured frequency zone is defined based on measured physiological signals associated with the intrinsic contraction of a muscle.

A stimulation frequency zone is defined based on the stimulation frequency of an electrical pulse generator configured to stimulate the muscle or a nerve associated with the muscle at a different frequency than the intrinsic frequency.

The muscle or a nerve associated with the muscle is stimulated, and the impact of the stimulation is determined by comparing the measured power of a set of signals falling within the measured frequency zone with the measured power of the set of falling signals. within the stimulation rate zone.

In variations of this embodiment, the physiological signal is a signal representing respiration.

In another embodiment, the muscle is a diaphragm muscle.

In another embodiment, the measured frequency zone comprises a range of frequencies immediate to the respiration rate of a patient.

[0033] In another embodiment according to the present disclosure, a system for electrical stimulation of a nerve or muscle includes an electrical pulse generator that supplies energy to stimulate a muscle at a first frequency, the first frequency different from a second frequency at which the muscle inherently contracts.

The system of this embodiment also has a sensor that detects physiological activity indicative of contraction of the stimulated muscle and an electronic memory for storing a collection of data generated by the sensor over a period of time.

A frequency analyzer analyzes the data collection to determine the power distribution across a frequency band for sensed physiological activity, and the system circuitry is able to compare the power of a band immediate to the stimulation frequency with that of the total power across the frequency band for detected physiological activity.

[0034] In another embodiment according to the present disclosure, a system for electrical stimulation of a nerve or muscle includes an electrical pulse generator that supplies energy to stimulate a muscle at a first frequency, the first frequency different from a second frequency at which the muscle inherently contracts.

The system of this embodiment also has a sensor that detects physiological activity indicative of contraction of the stimulated muscle and an electronic memory for storing a set of data generated by the detector over a period of time.

A frequency analyzer analyzes the data set to determine the power distribution across a frequency band for detected physiological activity and the system circuitry is able to compare the power of a band immediate to the stimulation frequency with the power total across the frequency band for detected physiological activity.

In this embodiment, the circuit is capable of increasing the energy delivered by the pulse generator if the ratio of the power of the immediate band to the stimulation frequency to the total power is below a threshold.

[0035] In another embodiment in accordance with the present disclosure, a system for treating disturbances

Respiratory ES 2 559 933 T3 includes an electrical pulse generator capable of providing electrical stimulation signals to a phrenic nerve or diaphragm of a patient with a predetermined signal frequency.

This embodiment has a breathing sensor capable of detecting a signal representative of the patient's breathing and a frequency comparator that compares the power density of the frequency distribution of the breathing signals detected over stimulation frequency bands with the power density of the breathing signals detected over a respiratory frequency band. The system includes a power adjustment circuit that adjusts the power of the electrical stimulation signals provided by the electrical impulse generation based on the comparison of the power densities. In various embodiments, the power setting circuit could adjust the stimulation current, the stimulation voltage, the frequency of the pulses in a pulse train, a pulse duration of the pulses in a pulse train, or others. parameters.

[0036] In another embodiment according to the present disclosure, a phrenic nerve or diaphragm is stimulated at a rate below an intrinsic rate of respiration.

Stimulation is delivered at an intensity sufficient to pace breathing while the intrinsic drive to breathe remains intact.

In some embodiments, the respiratory drive manifests as a 2: 1 pacing, in others as a spontaneous breathing of an unstimulated lung, in others as smaller breaths interspersed between paced breaths, and in others as a periodic activation of muscles. of the respiratory tract.

[0037] In another embodiment in accordance with the present disclosure, respiratory disorders are treated by detecting a signal representative of a patient's respiration and by frequency analysis of the signal representative of respiration over a range of frequencies. that matches the breath.

This embodiment includes the steps of determining an intrinsic respiration rate or rate and stimulating a patient's hemidiaphragm at a rate other than the intrinsic respiration rate.

In this embodiment, a frequency analysis of the respiratory signal is performed during stimulation.

A capture index is determined by dividing the power distribution in a frequency range immediate to the stimulation rate to the power of the frequency range according to respiration to determine a capture index.

The stimulation parameters are modified based on the calculated capture index.

In some embodiments, the representative signal for respiration may be transthoracic impedance. In some embodiments, the stimulation rate is less than the intrinsic respiration rate.

In another embodiment in accordance with the present disclosure, respiratory disorders are treated by detecting a signal representative of a patient's respiration and performing a frequency analysis of the signal representative of respiration over a range of frequencies according to the breath.

This embodiment includes the steps of determining an intrinsic respiration rate or rate and stimulating a patient's hemidiaphragm at a rate other than the intrinsic respiration rate.

In this embodiment, a rate analysis of the respiratory signal is performed during stimulation.

A capture index is determined by dividing the power distribution in an immediate frequency range from the stimulation frequency to the power of the frequency range according to respiration to determine a capture index.

The stimulation parameters are modified based on the calculated capture index.

In some embodiments, the stimulation power is increased if the capture rate is below a certain threshold.

In some embodiments, the stimulation power is lowered if the capture rate is above a certain threshold.

Brief description of the drawings

[0039]

Figure 1 is a waveform of respiration data from an untreated patient.

Figure 2 is a waveform of the respiration of the patient of Figure 1 during treatment in accordance with embodiments of the invention.

Figure 3 is a spectral graph of the respiration data presented in Figure 1.

Figure 4 is a spectral graph of the respiration data presented in Figure 2.

Figure 5 is a graph of experimental data on respiratory disorders in a patient treated in accordance with embodiments of the invention.

ES 2 559 933 T3

Figure 6 is a waveform of respiration and pacing data associated with treating a patient in accordance with embodiments of the invention.

Figure 7 is a control flow chart in accordance with embodiments of the invention.

Figure 8 is a schematic view of a patient and treatment device in accordance with embodiments of the invention.

Detailed description

[0040] Figure 1 is a waveform of the respiration data of an untreated patient.

The waveform represents 60 seconds of data acquired during the time the patient was asleep.

Trace 101 represents normal (at rest) breathing for this patient.

Trace 101 was acquired at 10:42 p.m. just before the patient received treatment.

Trace 101 represents the flow of air in and out of the patient's lungs monitored with a flow meter (thermal sensor).

At that time the patient is not showing periodic breathing or apneas and the patient is breathing regularly at 24 breaths per minute (0.4 Hz).

[0041] Figure 2 is a waveform of the respiration of the patient of Figure 1 during treatment according to embodiments of the invention.

This waveform also represents 60 seconds of data acquired when the patient was asleep.

The patient has ACS and was treated using transvenous stimulation of the right phrenic nerve.

Trace 103 represents stimulation pulse trains.

The pacing pulse trains were delivered at a constant rate of 18 per minute (0.3 Hz pulse frequency), in this case below the patient's native respiration rate of 24 breaths per minute. Each pulse train lasts 1.67 seconds.

During the application of the train of impulses, the right phrenic nerve was stimulated and the muscles of the right hemidiaphragm contracted (stimulation phase 104).

Each pulse train is followed by the relaxation phase 105 which also lasts 1.67 seconds in this example. During the relaxation phase, the phrenic nerve is not stimulated.

The stimulation model 103 therefore represents 50% of the work cycle stimulation (50% inspiration - 50% expiration).

Other duty cycles or ratios can be used to achieve the objectives of the invention.

Each pulse train 104 is composed of series of individual pulses (not shown) supplied by a pulse generator.

The pulse generator can be external or implanted.

In this example the pulses last 150ps (microseconds) and are applied at a frequency of 20 Hz.

The pulse generator in this example applied pulses in a controlled preprogrammed manner to achieve a preprogrammed maximum current amplitude of 4.9 mA.

To increase patient comfort, the electrical current amplitude of the pulses in the pulse train can be gradually increased, held constant, and then gradually decreased in the same pulse train.

Other forms of pulse train can be used to elicit the desired contraction and relaxation responses of the diaphragm muscle without departing from the scope of this disclosure.

[0043] Trace 102 represents the patient's respiration during stimulation therapy.

Similar to trace 101 in Figure 1, it consists of individual breaths monitored by a flow meter.

Each breath consists of an inspiration phase 106 and an expiration phase 107.

It can be seen that there are mostly large breaths 108 and some smaller breaths 109.

Large breaths are synchronized in phase with the pacing pulses 103 and the patient's breath can be said to be in time with the pacing.

These large breaths appear at the same rate as the pacing pulse train rate of 18 per min (0.3 Hz).

[0044] When only large breaths appear in series, they indicate 1: 1 pacing (one breath for each stimulation impulse train).

When small breaths appear, they can appear in several ways.

In some cases, small breaths are sandwiched between large breaths at a rate that generally corresponds to the spontaneous breathing rate of 24 breaths per minute (0.4 Hz).

In some cases, small breaths appear only occasionally as required by metabolic demand.

In other cases, periods of interleaved breaths correspond to a 2: 1 pacing (two breaths for each train of pacing pulses).

ES 2 559 933 T3

In any case, the existence of small, spontaneous or 2: 1 paced breaths supports the inference that while breathing is paced in this way, the respiratory control center remains active.

[0045] While the breaths in trace 102 appear at the same base rate as pacing 103, they are not necessarily exactly synchronized to pacing.

Inspiration 106 may start at a different delay time after the start of the pacing pulse trains 104 and may even under certain circumstances precede the appearance of the corresponding pacing pulse train.

In the embodiment shown in figure 2, the stimulation is applied at a rate somewhat lower than the spontaneous breathing rate at rest (in this example: 18 vs. 24 / min).

Stimulation can accompany spontaneous breathing but does not replace it.

Pacing is evident due to the variable time delay and phase angle between the pacing pulse trains and the patient's inspiration effort.

Further evidence of pacing is the periodic appearance of 2: 1 and 1: 1 pacing or other small breaths that are not paced, showing that the respiratory drive is not suppressed, but paced.

[0047] Figure 3 is a spectral graph of the respiration data presented in Figure 1.

Periodic shape waves such as the respiratory shape waves illustrated by Figure 1 can be analyzed using various frequency domain methods, the most common of which is spectral analysis.

[0048] Figure 3 shows the patient's normal spontaneous respiration spectrum illustrated by trace 101 in Figure 1.

Such a spectrum can be obtained by performing the Fast Fourier Transform (FFT) in 2-3 minutes of the digitally obtained respiratory signal data (in this example, the air flow).

The spectrum can be a power spectrum, a power density spectrum, or a magnitude spectrum.

The power spectrum can also be considered to provide which frequencies contribute the most to the variance of the signal.

The larger the amplitude, the higher the variance.

This is a significant broad definition of power spectrum. It is understood that many numerical methods exist for calculating the frequency distribution of periodic signals, and they are all contemplated here.

The spectrum can be calculated for the range of natural respiratory frequencies which are generally between 0 and 1.0 Hz.

In the disclosed embodiment, the frequency range of about 0.1 to 0.5 Hz was discovered. The selected range is called the respiratory frequency band (RFB) for the purpose of this disclosure. Other frequency ranges could be selected, and the selection of this range is merely illustrative.

[0049] In Figure 3, RFB 204 is indicated by the square frame that includes all respiratory rates important to determine pacing efficiency in this example.

The fact that the patient's natural respiratory rate has a maximum of 0.4 Hz 202 can be expected from the respiration trace 101 in Figure 1.

Another important frequency band for determining pacing is called the stimulation frequency band (SFB).

It is represented by the square narrow band 203 and in this case it is centered on the frequency at which the stimulation pulse trains are applied as in trace 103 in Figure 2.

The spectral graph of Figure 3 corresponds to the period in which the stimulation therapy is not applied.

Power in the SFB band is low if the center frequency, in this case 0.3 Hz, is significantly different from the dominant respiratory rate 202.

[0051] The efficiency of respiratory pacing can be determined by comparing the spectral power of the SFB band with the total spectral power or with the spectral power of other frequency bands, for example.

The inventors have found it useful to designate a capture index (CI) as a measure of the effectiveness of the treatment.

The capture index is calculated by dividing the spectral power of the SFB by the spectral power of the RFB.

The capture index represents the fraction of the total spectral graph of respiration that falls in the narrow band immediately at the pacing rate.

The value of the capture index can be expected to increase proportionally with the pacing of respiration through stimulation.

Thus, the pacing rate becomes the dominant rate of the respiratory signal as the pace increases.

There are many numerical calculations that can be used to calculate the catch rate.

For example, the inventors have used the following methodology, among others.

ES 2 559 933 T3

[0052] A spectral graph is a graphical technique for examining cyclic structure in the frequency domain. Strictly defined is a smoothed Fourier transform of the autocovariance function.

Frequency is measured in cycles per unit of time.

The spectral graph is displayed with a smooth vertical axis of variance (power) and a horizontal frequency axis (cycles per observation).

Computations to generate the smoothed variances can be included and are not discussed further here. Spectral graphs are a fundamental technique in time series frequency analysis and are familiar to those of skill in the art. Spectral graphs can be used to determine how many cyclic components are in a cyclic waveform, whether or not there is a dominant cycle frequency, and if there is, what is the dominant cycle frequency.

For the purposes of this disclosure, the degree of dominance of the dominant frequency (pacing) refers to the capture rate.

[0054] In the example shown in figure 3, the IC (capture index) was computed as: (power in the stimulation frequency range +/- 0.0183 Hz) / (power from 0.1 to 0, 5 Hz).

The width of the Numerator in this example is 0.0366 Hz.

It is calculated from the pacing rate value rounded to the closest FFT class with the Numerator width of 3 classes below and 3 classes above (6 classes in total).

The resolution in the frequency domain in this example is 0.0061 Hz / bin.

This is called the class width. The choice of +/- 3 classes was to account for something or the class distribution seen when the pacing rate is not an exact FFT class rate and for some natural variance in the rate of the paced breath.

The FFT intervals are set by the data sampling rate (400 samples / sec) and the FFT length (216 = 65,636).

[0055] The data sampling rate was 400 samples / sec.

It is understood that other lower sample rates, for example 20 samples / second, may also be suitable for calculations or other embedded software applications.

The length of the record was 65,536 samples (this corresponds to 164.84 seconds).

If a lower data collection sample rate is used, significantly fewer samples are needed, but less than, for example, 1,024 or 2,048 is unlikely.

Other record lengths, such as for example 2,3 and 5 minutes were also used with success.

In this example, the capture index is calculated as the ratio of the sum of the magnitudes of a frequency band ± 3 classes wide (0.0366 Hz), centered around the known stimulation index and the sum of the magnitudes (or power) 0.1 to 0.5 Hz.

Other window widths for the SFB and RFB could be used in different embodiments, and the SFB does not need to be centered on the pacing rate.

Common to all embodiments, the stimulation frequency band SFB is narrower than the respiratory frequency band (RFB), such as for example <10% RFB and included in the RFB.

[0056] Figure 4 is a spectral graph of the respiration data presented in Figure 2.

Stimulation has been turned on in this example and applied at the programmed frequency of 0.3 Hz as illustrated by trace 103 in Figure 2.

It can be seen that the maximum value 302 of the power spectrum of the respiratory signal (air flow) is at the frequency 0.3 Hz which corresponds exactly to the stimulation frequency.

Natural breathing power at 0.4 Hz is reflected by a lower maximum value 301.

It is appreciated that the presence of the natural respiration frequency power in the spectrum may vary depending on the stimulation force and the intrinsic respiratory drive of the patient.

The presence of noticeable respiratory activity at the natural respiratory rate suggests that the respiratory drive is paced but still active, that is, it is not suppressed or dormant.

The power of the SFB 303 in this example represents a larger fraction of the RFB 304 than in the example illustrated by Figure 3.

Therefore, the capture ratio (SFB / RFB) can be expected to increase significantly as well.

In fact, calculations show that the catch rate increased in this example from 0.24 (Figure 3) to 0.46 (Figure 4).

Different calculation methods can result in different numbers, but the capture rate increases significantly when the patient's breathing is paced.

[0058] Figure 5 is a graph of experimental data related to respiratory disorders in a patient treated in accordance with embodiments of the invention.

Figure 5 illustrates the practical importance of catch classification in order to restore normal respiration in the adjustment of periodic respiration.

The severity of periodic breathing is commonly characterized by the apnea hypopnea index (AHI).

The AHI is the sum total of respiratory events (apneas and hypopneas) that occur in one hour.

ES 2 559 933 T3

An AHI> 15 is considered significant and an AHI> 30 serious and very dangerous.

There are known standard clinical methods for calculating AHI during sleep studies using polysomnography (PSG).

PSG is a diagnostic test during which various physiological variables are measured and recorded during sleep.

The graph in Figure 5 was obtained by investigators using PSG in a patient suffering from severe periodic breathing while undergoing stimulation treatment according to embodiments of the invention.

The patient's AHI is expressed on the Y-axis as a function of the capture rate on the X-axis.

During the experiment, the stimulation energy (in this case, stimulation current) was varied generating different levels of diaphragmatic activation and, as a result, different levels of pacing. The capture rate was later calculated using a digitized record of the sleep study and a methodology equivalent to those described in this disclosure.

[0059] It can be seen that during periods of time when the capture rate was higher, the AHI was reduced.

Stimulation that resulted in capture rates greater than 0.5, as calculated in this example, virtually completely eliminated periodic respiration.

[0060] It is understood that the use of FFT and respiratory spectrum calculation is not the only way to implement capture rate calculation.

For example, the respiratory waveform can be processed and displayed as a series of numbers that correspond to the duration of the breaths.

A series representing the last 3-4 minutes of respiratory data may consist of, for example, 60-80 breath durations.

A breath durations histogram can then be constructed that represents the frequency distribution of the breath durations.

If intrinsic respiration is paced to a pacing rate, the incidence rate of breath durations that corresponds to the wavelength of this rate will increase.

For example, if the stimulation rate is 20 / minute, the duration of the breath is 3 seconds.

As capture and pacing increase, breaths of approximately 3 seconds in length will occur more and more frequently.

To compensate for natural variability, breaths that last for example between 2.84 and 318 seconds can be included in the numerator of the capture rate calculation.

This range of breath durations corresponds to the +/- 0.0183 Hz frequency band used in the spectrum-based example described above.

[0061] The capture index in this method is computed as: (sum of occupations of durations of breaths in the stimulation frequency range +/- selected band) / (sum of occupations of all durations of breaths in the respiratory range) .

The respiratory range of the breath durations can be for example 2 to 10 sec.

The breath duration range from 2 to 10 seconds corresponds to the respiratory rate 0.1 to 0.5 Hz.

[0062] The histogram-based catch verification method is mathematically different from the spectral-based method, but similar in principle.

It is based on the assumption that when pacing paces the breath, the durations of breaths of the respiratory signal in the immediate range of the pulse rate will occur at a significantly higher rate than those in other respiratory rate bands.

Therefore the capture index (CI) is still calculated as the ratio of the variance of the respiratory signal that falls in the narrow band centered on the pacing rate to the total variance in the broader respiratory rate band.

The inventors have shown in patients that the capture index value calculated using the histogram method increased proportionally to the pacing of breath by stimulation, which allowed a guided treatment and closely correlated with the capture calculated using the spectrum method.

It is appreciated that other catch rate calculation methods based on similar principles are possible and will occur to those skilled in the art upon reading this disclosure.

[0063] Figure 6 is a waveform of respiration and pacing data associated with treating a patient in accordance with embodiments of the invention.

The upper line 501 represents a patient's breath (air flow).

The X-axis represents 16 minutes of logged data and individual breaths are compressed compared to Figures 1 and 2 which showed only one minute of data.

During this treatment period the stimulation energy (current) was first gradually reduced and then completely deactivated.

The lower trace 503 represents the electrical stimulation current.

During the first 3 minutes the current was kept constant at approximately 5 mA.

ES 2 559 933 T3

Respiration 501 was sufficiently paced and periodic respiration was not present.

During the 3-10 minute period of the recorded segment the current 503 was gradually reduced.

It can be seen that periodic breathing is no longer controlled, and after 10 minutes the alternating apneas 504 and hyperpneas 503 returned, indicating the typical periodic breathing pattern known as Cheyne-Stokes respiration.

[0064] Figure 7 is a control flow chart in accordance with embodiments of the present disclosure.

Figure 7 illustrates a potential method and algorithm that uses catch classification to implement and improve treatment for respiratory disorders.

The proposed capture rate methodology is understood to have broad implications for respiratory treatments using phrenic nerve or diaphragm stimulation and potentially for mechanical ventilation.

[0065] An embodiment in accordance with the present disclosure employs an index-type capture calculation input into an algorithm in an implantable pulse generator (IPG) microprocessor that is capable of adjusting phrenic nerve stimulation energy in response to calculated catch rate.

The capture rate can be calculated based on 2-3 minutes of the history of a respiratory signal, for example in transthoracic impedance, and the stimulation parameters can be automatically adjusted based on the calculated capture rate.

The algorithm introduced in this embodiment is capable of determining the intrinsic resting respiratory rate of patient 601.

This rate can be, for example, between 6 and 40 breaths / minute, but in a predicted population of patients with periodic breathing, between 12 and 30 breaths / minute is possible.

The algorithm is capable of applying trains of stimulation pulses to the patient's phrenic nerve at a set rate that in some embodiments is somewhat lower than the intrinsic rate 602.

It can be, for example, 2-4 breaths lower than the intrinsic index.

Therefore, if the algorithm has determined that the patient was breathing at 20 breaths per minute, the pacing rate can be automatically set to 16 breaths per minute, for example.

After the entered software has collected sufficient respiratory signal information (this may be for example 3 minutes of digitized data at 20 samples per second), a capture rate 603 can be calculated.

[0067] Since physiological conditions such as the patient's posture, sleep state, diaphragm fatigue, and others can influence the response to pacing, it is to be expected that the capture rate will not be constant over time.

There may be a programmed assigned Capture Index value that indicates the desired timing.

This value can be in the range of 0.2 to 0.8 for example.

According to the known technique of feedback control engineering, an introduced algorithm can compare the actual capture rate with the target 604 and increase the stimulation energy if the capture rate is below the target or decrease it if it is above it. 605.

The stimulation energy can be adjusted by varying the delivered current, voltage, frequency, or pulse duration.

For example, the stimulation current can be increased or decreased in the range between 1 and 10 mA in appropriate steps.

Alternatively, the duration of the stimulation pulse can be increased or decreased in the range of 100 to 500 ps in suitable steps, for example.

A feedback control algorithm can be implemented in the introduced IPG software, such as a PI or PID regulator known in the field of control engineering.

In the embodiment used by the inventors to collect experimental data, the stimulation current was manipulated in 0.1 to 0.5 mA steps to achieve the desired capture rate.

[0068] Figure 8 is a schematic view of a patient and treatment device in accordance with embodiments of the present disclosure.

An implantable pulse generator 702 is programmed to generate stimulation pulse trains 703 at a fixed rate.

A right phrenic nerve from patient 705 innervates the right hemidiaphragm 706.

Stimulation pulse trains 703 are transmitted to right phrenic nerve 705 through electrode wire 704.

[0069] Either the phrenic nerve or the hemidiaphragm itself could be stimulated.

One or more electrodes could be placed on the diaphragm, adjacent to the nerve (eg, cuff electrode), intravenously close to the nerve, or at any other suitable location to provide appropriate stimulation.

The electrode (s) could be connected to an electrical pulse generator using wired or wireless technology.

ES 2 559 933 T3

The pulse generator could be implanted in the patient or be located externally.

[0070] The right phrenic nerve 705 conducts stimulation to the right hemidiaphragm 706, which responds with downward movement 707.

The downward motion 707 of the diaphragm inflates the lungs and activates stretch receptors within the thoracic cavity.

Periodic stretching is believed to generate a periodic regular rhythm of neuronal feedback inputs 708 to the brain 701.

Respiratory neurons in the brain are paced by neuronal input 708.

The stimulation force 703 elicits a proportional response from the stimulated hemidiaphragm 706.

Increased diaphragmatic movement consequently results in increased regular periodic neuronal input 708 to the brain 701.

[0071] When the signals reach the necessary force, respiratory pacing is present and the desired fixed and repetitive coupling is temporarily established between stimulation, mechanical inflation of the lungs and inspiratory neuronal activity in the respiratory control center of the brain.

Pacing can occur at a 1: 1 ratio (a mechanical inflation for a neural respiratory effort), but other integral ratios can be seen, as well as occasional aperiodic and chaotic behavior in the transition between pacing patterns with a different integral ratio. .

The 2: 1 ratio of two paced inflates for neuronal respiratory effort is commonly seen in conjunction with the 1: 1 ratio.

[0072] The brain responds to timing with the regular periodic sequence of respiratory impulse 709 that is delivered to the respiratory muscles of the diaphragm through both the right 705 and left 712 phrenic nerves as well as through the control nerves of the airways. airway 710 to the airway muscles resulting in the desired dilation of the airway 711.

The hemidiaphragm 715 is innervated by the left phrenic nerve 712 which is not stimulated by the IPG and therefore may occasionally exhibit independent behavior in response to signals coming from the brain 701 and is not directly affected by the IPG 702.

The muscle groups of the right and left hemidiaphragms are known to be innervated separately by the right and left phrenic nerves and move independently in response to signals from these nerves.

The synchronized respiratory activity of the unstimulated hemidiaphragm is an indication of pacing rather than the rhythm of respiration.

The IPG 702 can be equipped with additional cables 713 and means to measure respiration such as through transthoracic impedance detection 714.

Software embedded in the IPG's programmable logic can respond to changes in respiration by adjusting the rhythm of the 703 stimulation pulse train.

Respiratory detection 714 can also be used by IPG logic to adjust and change the rate of pacing pulse trains 703 depending on the intrinsic respiratory rate detected 709.

[0074] One skilled in the art will appreciate that the present disclosure can be practiced with embodiments other than those described.

The described embodiments are presented for the purpose of illustration and not of limitation.

The invention is defined only by the claims that follow.