Spirometric device

<td colspan="2">SWEDEN</td><td rowspan="2">(12) EXPLANATORY NOTES O (51) International class</td><td rowspan="2">IBl (2i)</td><td rowspan="2">7610356-3 A 61 B 5 / Ö8</td>

<td> (19</td><td>) SE</td>

<td></td><td>ww UAif</td><td>(44) Application submitted and submitted 81 —05—1 8</td><td>(11) Publishing</td><td> 418 244</td>

<td></td><td>tSa</td><td>publication published</td><td>number</td><td></td>

(41) (22) (24)

PATENTVERKET «“&gt;

Application publicly available 77-03-19 The patent application was received on 76-09-17

Running day '

The number of the main application

Application received as · K Swedish patent application (86) International filing date (86) Filing date for European patent application (30) Priority information

75-09-18 FR 75 28561 □ completed international patent application with number

In converted European patent application with number

KB 3-A4 according to SIS 61 30 13

71Applicant: SYNTHELABO SA, PARIS, FR

72) Inventors: RPC Cavallo, Bourg-la-Reine and A Le Maitre, St-Maur

74) 0mbud: Delmar

54) Name: Spirometry device

7610356-3

The present invention relates to a spirometry device connected to a patient's respiratory tract system and comprising an outflow sensor device intended to be mounted in the exhalation branch, the signals emitted by a pressure transducer arranged in the airway system and the signals emitted by the outflow sensor device being fed to a calculation and arranged to integrate the outflow signals and correct them depending on the pressure variations.

A spirometer measures the volume or flow of exhaled gas, which volume is normally referred to as "the flow volume ooh mainly corresponds to the respiration capacity, which is hereinafter referred to as the effective" capacity because it corresponds to the amount of gas that the lungs can actually exhale and is obviously less than actual capacity, as the lungs are not completely emptied during each respiratory cycle up to the point of collapse.

To date, a number of different methods of artificial respiration have been used without, however, being entirely satisfactory. As examples of the state of the art, the following patents can be mentioned:

In accordance with U.S. Pat. No. 5,006,356, an attempt is made to compensate for the pressure loss in the spirometer which counteracts the patient's

7610356-3 exhalation by means of a piston, which is regulated in order to maintain a constant pressure in the chamber of the spirometer.

Many devices are further known for measuring volumes of flows and with pressure correction by applying Boyle-Mariotte's law. Examples which may be mentioned in particular are U.S. Pat. Nos. 5,759,249 and 5,799,149, which relate to a breath analyzer having a mass spectrometer and a metabolic analyzer, respectively, in which a unit of calculation introduces a pressure correction by applying Boyle-Mariote's law.

That the pressure is regulated and fed is further known from U.S. Patent No. 5,414,896.

Virtually all known devices operate with an approximate pressure measurement and do not take into account variations in this pressure, which can often be significant during one and the same cycle.

The object of the present invention is to provide a device without the above-mentioned disadvantages.

This is achieved by means of the device mentioned at the outset, which has obtained the characteristics stated in the claims.

Thanks to the device according to the present invention, the measurement becomes more accurate at the same time as both the maximum and minimum pressure are taken into account, as well as the over- or under-pressures that occur during the breathing cycle.

The solution according to the present invention is furthermore very uncomplicated and, compared with other solutions such as, for example, the mechanical solution according to U.S. Pat. The invention is described in more detail below with reference to the accompanying drawing, in which Fig. 1 shows a conventional system for artificial respiration, Figs. 2 and 4 show schematic diagrams of two versions of the measuring and calculation unit M in Figs. 1 and Fig. 5 shows a block diagram of the embodiment according to Fig. 2.

Corresponding elements in the various figures have been provided with the same reference numerals. In these figures, it is assumed that negative pressure prevails. If there is no negative pressure, it is sufficient to replace the index 2 with the index 0.

With reference to Fig. 1, a conventional artificial respiration apparatus mainly comprises a Y-shaped pipeline, one branch of which is connected to the patient P and the other branch of which is connected to a source of pressurized gas (for example a tube or a container

7610356-3 and a compressor) via a valve EV1, which is normally electric in newer appliances, and whose third branch is connected to the surrounding atmosphere via a valve EV2, which is also electric in most cases and which can possibly be followed by a spirometer S.

It is known that when the patient needs artificial respiration, ie. if it is thought that he is completely or partially unable to spontaneously ensure his own breathing, the patient must be exposed to a certain overpressure compared to atmospheric pressure during inhalation (EV2 is open and EV1 is open), which is higher the more difficult the patient is to breathe. The inhaled volume thus has a pressure which differs from the outside atmospheric pressure, which corresponds to a gas mass which is greater than that of an identical volume at atmospheric pressure. Upon exhalation (EV1 is closed and EV2 is open), the pressure is further reset progressively and while the air is diverted via the spirometer, its pressure varies, ie. the initial overpressure decreases until it is zero, so the spirometer measures a volume under a pressure that varies during the measurement.

The effect of the overpressure on the volume can be neglected as the overpressure itself is negligible, but normally the overpressure must be taken into account.

The present invention offers a solution to this problem by means of a sensor K, which measures the pressure in the Y-line and which transmits data to a measuring and calculation unit M, which also receives data from the spirometer S and directly corrects its measurements in the desired manner. Thus, if the index 0 is attributed to the atmospheric conditions and the index 1 the maximum overpressure conditions (which may vary from one breathing cycle to another), M no longer integrates the volumes passing into the spirometer, but these volumes are corrected to atmospheric pressure.

As will be apparent from the two embodiments described below, there are two solutions, namely an approximate one, where the exhalation volume is integrated for multiplication with the pressure variation P1 - Po Γ ⁰i.e. —--- L dV, and a more precise, where the flow corrected for atmospheric pressure is integrated in each moment ie. p_Q dV, where P is the instantaneous pressure measured by the sensor K, the instantaneous elemental volume dV or the instantaneous outflow dV / dt being measured by S. Boyle-Mariotte's law, which applies to Ideal gases or to real gases at small pressure variations, which is usually the case within the spirometry, applies in the present case, but it would

7610356-3 be possible to integrate using another law in extreme cases ie. if the pressure variations were to be very strong. This is possible, for example, when people with diving disease are revived in a pressure chamber.

Account must also be taken of the fact that modern appliances not only work under overpressure when inhaling but_OP1 - pp also under negative pressure on exhalation (indes 2), why -5 ^ --- £ dV / * 2 p • '' 1 or J -pj dV is thus integrated in M. Corrections can also be made if the ^ spirometer itself causes a pressure loss that dampens the pressure variations over time.

Furthermore, it should be emphasized that the volume of the lines is generally not negligible and that this must be taken into account as the most important information for the doctor is the amount of gas emitted by the lungs. This information can be obtained without correction at the mouth of the sick person by measuring the outflow and pressure, but this measuring point is unsuitable due to the fouling caused by the condensation of water vapor or the patient's excretions which make the measurements incorrect. This explains why today one seeks to place the sensors, especially those intended for the flow, far enough away from the sick person's mouth, leading to a non-negligible increase in the length and volume of the wires.

Accordingly, it is convenient to reduce the volume at atmospheric pressure corresponding to the gas mass which itself corresponds to the difference between the gas masses in the line between P1 and Po.

The outflow sensor S (Figs. 1 and 2) transmits flow data in liters to the unit M during the exhalation part of each cycle, for example by transmitting a number of pulses proportional to the flow rate in the sensor and thereby to the flow. This can be obtained, for example, by means of a capacitive type turbine sensor.

The integration with respect to the time in the integrator 2 gives the so-called “flow volume” for each cycle. In the integrator 5, the so-called the minute volume ”. These data correspond to the total volume delivered by the respirator (at varying pressures during a cycle).

If the capacity of the pipes is Co independent of the pressure P, a volume C - Co - Co flows through the outflow sensor in addition to the breathing volume, ie. 4^c “ ^{Co om} there is nothing

CoP1 CoP2 negative pressure on exhalation ooh Zi C = - ie.

Λ ^{c om} the case is the opposite.

7610356-3 (The volume Co of the gas in the lines is in fact Co when changing from pressure P to pressure Po.)

The pressure sensor K supplies the unit M at 4 with a pressure measurement whose maximum P1 is marked at 5 and minimum (Po or P2 depending on whether negative pressure prevails or not) is marked at 6, while the difference is obtained at 7 ·

This pressure change is transmitted to the correction devices 8 and which receive the "minute volume"¹¹ and the flow volume from 3 and 2, respectively, and emits the corrected “minute volume” and the “flow volume”, respectively, with regard to the pressure variations and the correction factor ÄC. The indication takes place at 10 and 11, respectively.

As an example, this device has been simplified because x P1 - P2 it does not perform the integration dV but --- gg— = J dV = ~ ^ pÖ '~ ^ 1 ⁹ ^1 ^RE^^{the journey}^^{erar en} volume that passes into the sensor between 1 and 2, ie. during the exhalation period. If, on the other hand (Fig. 4), the instantaneous pressure P is transmitted from 4 to an instantaneous correction device 12 upstream of the integrators 2 and 3,_r2 p the gration as above and the integrators 2 and 3 emit / -pg dV, while the correction devices 8 and 9 in this case only give the correction applicable to Δ C.

With reference to Fig. 3, it will be explained how circuits for performing these operations can be constructed. If △ C is the correction for the flow volume, the correction η Δ C must be applied to the minute volume, where n is the number of breathing cycles per minute. The measurement of n is not shown in Fig. 2 or 4, but an embodiment of this will be given in Fig. 4.

Several solutions are in fact possible: n is the number of pressure cycles and can thus be counted in K (each time P = Po if no negative pressure prevails or half of the number of times when P = Po or P2 has a minimum) or in S each time the outflow becomes zero (inhalation), which solution has the advantage that it is easy to apply. It is also possible to calculate the closures and openings of EV1 or EV2 or otherwise to calculate the ratio between the minute volume "and the flow volume" etc.

Fig. 3 shows an embodiment of the principle diagram according to Fig. 2.

7610356-3

The majority of circuit elements in Fig. 5 are identical and the change should be apparent to those skilled in the art. The spirometer consists of a turbine 20, which is followed by a capacitive sensor 21 and a voltage comparator 22. The shape of the transmitted signals has been shown schematically in the figure.

The comparator 22 acts on the one hand on a monostable circuit 23 and on the other hand on a detector 24, which comprises an RC circuit which is pulse blocked for zero reset. The signals from the monostable circuit 2J are transmitted on the one hand to an RC minute surface integrator 25j which is followed by a direct current amplifier 26, which thus emits an analog signal proportional to the minute volume. The signals from the monostable circuit 2? is also transmitted to an RC cycle surface integrator 27 y which is likewise followed by a direct current amplifier 28, which emits an analog signal proportional to the flow volume. The one cycle operating integrator 27 is further actuated by the detector 24, which resets the integrator 27 to zero for each cycle and triggers a constant amplitude type monostable circuit 29, which is followed by a one minute operating surface integrator JO, which consequently emits one against n, ie. respiratory rate per minute proportional signal.

The analog pressure sensor K is followed by a direct current amplifier J1, which in turn is followed by a filter J2, which separates the peaks P1, and a filter JJ, which separates the valleys (or the negative peaks) P2 (or Po).

P2 is inverted in an inverter J4, whereby it becomes possible on the one hand to form P1 - P2 in an amplifier J5 of the signals from the units J2 and J4 and on the other hand to form n (P1 - P2) in an amplifier j6 of the signals from the unit JO . The amplifiers J5 Q and 36 thus emit proportional, analog signals against Δ and Δ01 min, respectively. The signals from the summing amplifiers 35 and 36 are adjusted by means of potentiometers 37 and 38, respectively, to calculate the actual volume of the lines, which can obviously vary depending on the application conditions. Thus, DC amplifiers 39 and 40, respectively, will receive analog signals corresponding to Δ and A Cl min, respectively, from potentiometers 37 and 38, respectively, and analog signals corresponding to v | and V ^ mn from amplifiers 28 and 26, respectively. Amplifiers 39 and 40 correct the volumes and emit corrected volume signals C1 and Vlmn-A Clmn, respectively. Amplifiers 39 and 40 are followed by output amplifiers 41 and 42, respectively, and the indication is obtained by means of a suitable means either in analog form such as, for example, by graded galvanometers 4j and 44 or in digital form by, for example, so-called nixie "indicators.

In the block diagram of Fig. J, only the correction of △ C has been shown for simplicity, since the correction for Δρ is negligible compared to the correction of Δ C in many cases.