The present invention relates in general to non-invasive mechanical ventilation systems which are used for 1) generating continuous positive pressure, also called Continuous Positive Airway Pressure - CPAP, in order to increase the residual functional capacity of the patient with alveolar derecruitment, counterbalance a residual pressure in the alveoli, also called PEEPi in patients with chronic obstructive pulmonary disease - COPD, reduce the afterload of the left ventricle in patients with acute heart failure, 2) supporting the activity of a patient's inspiratory muscles through an intermittent positive pressure (usually a flow- cycled pressure support also called Pressure Support Ventilation - PSV or BILEVEL, but also time-cycled pressure support, also called pressure controlled ventilation - APCV or biphasic pressure ventilation, also called PC-BILEVEL or PC-APRV, in order to deliver an adequate volume of gas to the lungs. Non-invasive mechanical ventilation with a helmet interface is widely used in critical areas as it has advantages over nasal or oral interfaces. Many patients who need non-invasive artificial ventilation prefer the helmet solution as this type of ventilation allows the patient to interact with people and the environment using their own voice and a field of view not obstructed by the interface. In addition, the helmet ventilation guarantees a good seal regardless of the physiognomy of the face, requires minimal contact with the skin and prevents skin lesions; the patient can also eat or drink through a straw inserted in the helmet through a sealing ring.

The main drawback of helmet ventilation is linked to the accuracy of the tidal volume measurements by the ventilator and the consequent reduction in the patient’s monitoring quality.

An object of the present invention is to provide a solution to allow a relatively precise estimate of the tidal volume in a helmet-based ventilation system through the ventilator itself and without the aid of other devices (e.g. inductance plethysmography, electrical impedance tomography, pneumotachograph to the patient's mouth). Therefore, the present invention relates to a method for estimating the tidal volume of a patient during non-invasive ventilation, by means of a ventilation system comprising a turbine driven ventilator provided with a flow sensor and a pressure sensor associated with the ventilator, a non-invasive ventilation helmet configured to be worn by a patient, said ventilation helmet comprising an inlet port and an outlet port, an inspiratory tube connecting an outlet of the ventilator to the inlet port of the ventilation helmet, and an intentional leak positioned on the helmet, wherein said method comprises measuring a flow of gas supplied by the ventilator, hereinafter machine flow, estimating a flow of gas through the intentional leak as a function of gas pressure within the helmet, hereinafter leak flow, estimating a respiratory flow of the patient based on said machine flow and leak flow, and determining a tidal volume of the patient by mathematical integration of the respiratory flow over time.

The inventors have discovered that the presence of an intentional leak allows the estimation of the tidal volume which is otherwise not measurable using a pneumotachograph (differential pressure transducer) of a compressed gas or turbine ventilator with a bi-tube circuit (separate inspiratory and expiratory branches). The measurement read by the ventilator in this case is the sum of the patient's tidal volume and the volume of the helmet. The measurement would be possible as done in some bench studies by inserting in the patient's mouth (obviously not feasible from the ethical and clinical point of view) a pneumotachometer provided with a mouthpiece or with other systems previously described but not through the ventilator itself.

As will be described, tidal volume estimation requires a turbine ventilator, a single circuit, and an intentional leak.

Further features and advantages of the invention will become apparent from the following detailed description of an embodiment of the invention, made with reference to the accompanying drawings, provided for illustrative and non-limiting purposes only, in which Figure 1 is a schematic representation of a mechanical ventilation system in which it is possible to measure the tidal volume.

With reference to Figure 1, a mechanical ventilation system comprises a turbine driven ventilator, indicated with reference numeral 10. Such a ventilator is conventionally configured to supply a gas flow to a patient P according to a predetermined assisted ventilation protocol, through an outlet 13, and is conventionally provided with a control unit for monitoring the ventilation operations. Associated with the ventilator 10 is a flow sensor 11 configured to provide a measurement of the gas flow supplied by the ventilator 10 to the outlet 13, defined in terms of units of liters per unit of time. A pressure sensor 12 is also associated with the ventilator 10 and is configured to provide a measurement of the pressure of the gas supplied by the ventilator 10 at the outlet 13.

The ventilation system further comprises a ventilation helmet 20 configured to be worn by a patient P. The ventilation helmet 20 is of a per se known type, and essentially comprises a cylinder of transparent material and a ring base configured to secure the helmet to the patient's body P. The ventilation helmet 20 conventionally comprises one or more ports or fittings with various functions, and in particular comprises an inlet port 21 and an outlet port 22. In particular, the helmet shown in Figure 1 is of the braceless type, which provides a base formed by an inflatable cushion which ensures the anchoring of the helmet to the patient's neck. However, the invention is not limited to the type of helmet shown, as it may also be applied to helmets with braces, which have a rigid base provided with braces that are fastened to the patient's armpits.

The ventilation system further comprises an inspiratory tube 30 connecting the outlet 13 of the ventilator 10 to the inlet port 21 of the ventilation helmet 20. The ventilation system is of the open circuit type with single tube and intentional leaks (also defined by the international nomenclature as “passive circuit” or “intentional leak circuit”), and therefore the outlet port 22 of the ventilation helmet 20 is not connected to the ventilator 10. An intentional leak 40 is in fact positioned on the outlet port 22, conventionally consisting of a device configured to facilitate the discharge to the environment of the carbon dioxide exhaled by the patient, thus avoiding the rebreathing phenomenon. The intentional leak may also be positioned in other parts of the helmet; in this case, the outlet 22 is closed. Unintentional leaks are also inevitably expected in the helmet, generally due to the imperfect sealing of the helmet in the areas in contact with the patient; however, these leaks are small and negligible compared to the intentional leak. The intentional leak may for example be made as a connector and provided with a hole having a diameter of less than 1 cm, for example about 5.5 mm, applied to the outlet port 22 or, as mentioned above, in other parts of the helmet.

A method for estimating a patient's tidal volume during a ventilation procedure using the system described above is now described. A similar method has hitherto been used to estimate tidal volume with nasal or facial interfaces.

As is known, the tidal volume is the quantity of air that is mobilized with each non-forced breath.

The flow, which hereinafter will be designated as machine flow ( Qmach ), is defined as the quantity of gas that goes from the ventilator to the patient. The respiratory flow Q _p of the patient (positive during the inhalation phase and negative during the exhalation phase) differs from the machine flow due to gas leaks (Qieak) ^' .

Qp Qmach ^~ Qieak (1) where Q _P is the flow inhaled and exhaled by the patient, Qmach is the flow supplied by the ventilator, Qi _ea k is the part of the flow supplied by the machine but not received by the patient (lost through leaks, mainly through the intentional leak 40).

The gas flow Qieak (leak flow) is then estimated through the intentional leak 40 as a function of gas pressure P _p inside the helmet, which hereinafter will be designated as a leak flow. The methods for determining the estimate of Qieak are known to ventilator manufacturers. Such an estimate is made using a mathematical function of pressure P _p applied inside the helmet. Once the flows Q _mach and Qi _eak have been determined, it is therefore possible to estimate the respiratory flow Q _p of the patient and consequently, through its mathematical integration overtime, the patient's tidal volume.

The inventors tested the configuration described above in bench studies and in healthy volunteers.

Regarding the bench study, a modified dummy head (Laerdal Medical AS, Stavanger, Norway) was connected to an active test lung (LS) (ASL 5000; Ingmar Medical, Pittsburgh, PA, USA). The ASL 5000 is a real-time digitally controlled breathing lung simulator, which allows you to create various types of spontaneous breathing patterns and different respiratory machine conditions (e.g. normal, restrictive or obstructive condition). Its operation is based on a direct-acting screw-driven piston, which moves inside a cylinder according to the equation of motion of an active respiratory mechanical system. Its settings during the bench study, using a single compartment model, were as follows:

(1) Normal condition: resistance 4 cmFUO/L/s. compliance 60 mL/cmFLO. inspiratory muscle pressure (Pmus) -5 cm FLO (half-sine waveform with 25% rise time, inspiratory hold of 5%, release time of 25%) and respiratory rate LS set at 15 breaths/min;

(2) Obstructive condition: resistance 15 cmLLO/L/s. compliance 80 mL/cmLLO. inspiratory muscle pressure (Pmus) -12 cm FLO (half-sine waveform with 20% rise time, inspiratory hold of 5%, release time of 30%) and respiratory rate LS set at 25 breaths/min;

(3) Restrictive condition: resistance 7.5 cmLLO/L/s. compliance 30 mL/cmFLO. inspiratory muscle pressure (Pmus) -12 cm FLO (half-sine waveform with 25% rise time, inspiratory hold of 5%, release time of 25%) and respiratory rate LS set at 30 breaths/min.

The dummy's head was connected to LS through the dummy's trachea after placing and attaching a helmet (Castar, Next Fntersurgical, small size, Mirandola, Italy). The inspiratory port of the helmet was connected to a turbine-driven ventilator (TDV) via a one-branch circuit while the expiratory port was closed with a lid having a 5.5 mm diameter hole to allow for intentional leak. Such a configuration was not provided with any type of inspiratory or expiratory valve or carbon dioxide discharges. During non-invasive pressure support ventilation by helmet (nHPSV) the ventilator was set to the shortest pressure rise time, and a cycling-off flow threshold of 25% in normal and restrictive conditions and 40% in obstructive condition. An inspiratory flow trigger initially set at 2 L/min was therefore always adjusted to the lowest value that did not result in self triggering. Inspiratory pressure above positive end expiratory pressure (PEEP) was set to achieve a tidal volume VT of about 300/500 ml in all conditions. Each condition was simulated at PEEP of 5, 8, 10 cmHiO. The PEEP of 12 cmHiO was used only in the restrictive condition. During the hHPSV, unintentional leaks were avoided by carefully placing the helmet on the dummy's neck. During the normal condition at CPAP of 8 cmH20, the reliability of the tidal volume estimate at different leak flows (30, 50 and 80 L/min) was also tested using a calibrated hole.

A study was also performed employing volunteers and using the same TDV configuration used in the bench study. A mouthpiece was inserted into the volunteer's mouth and connected to a pneumotachograph (P) (VT mobile FLUKE, Germany) to measure the subject's respiratory flow and VT. A nasal clip was placed on the volunteer's nostrils to prevent nose leakage. All the volunteers were instructed, using a metronome, to maintain an imposed respiratory rate of approximately 12/15 breaths per minute. During nHPSV, the ventilator was set to the shortest pressure rise time and a 25% cycling-off flow threshold. An inspiratory flow trigger initially set at 2 L/min was therefore always adjusted to the lowest value that did not result in self-triggering. During nHPSV, leaks were carefully avoided by choosing the correct helmet size and placing the helmet appropriately on the volunteer's neck.

Regarding bench measurements, data was collected by the TDV and LS software. Air flow, airway pressure (Paw), PEEP, as well as VT and respiratory rate provided to the dummy, were collected every 3 minutes. Differences in VT between TDV and LS were compared in the last 20 breaths of each test to ensure system stability between settings changes. When testing at different leak flows, each measurement was made 1 minute after the onset of the new leak to allow the TDV to adapt to the new setting. Self-triggering was determined as mechanical insufflation in the absence of inspiratory effort on the LS. Regarding the volunteer trials, each experimental condition was maintained for 3 minutes to ensure system stability between settings changes. Paw, PEEP obtained from the TDV and VT and respiratory rate of the TDV software and the pneumotachograph (P) were collected. The values of VT measured by the TDV and P were compared during the last 20 breaths of each test.

The main conclusions of the bench study are as follows: 1) There were no clinically relevant differences in VT between TDV and LS in normal and restrictive conditions at all simulated PEEP levels; 2) The difference between TDV and LS remained stable between the different leak flows tested; The main conclusions of the study on healthy volunteers are that although there were differences in VT on almost all volunteers, these differences were of little clinical relevance.