The main findings of this bench study are that, during helmet ventilation: (1) CO2 rebreathing can be monitored measuring the CO2 concentration at a “quiet” point inside the helmet or alternatively at the Y-piece, and (2) CO2 rebreathing is inversely related to the total flow passing through the helmet and directly related to the patient's CO2 production.
In our experience we observed that CO2 concentration is not homogeneous inside the helmet. When the sampling was performed between patient and ventilator, at airway opening or Y-piece, CO2 oscillated between low inspiratory and high expiratory levels. In contrast, CO2 was very stable when measured at a quiet point, not affected by flows to and from the patient; moreover, this value was really equivalent to the mean inhaled CO2. Therefore, when a “quiet” point inside the helmet can be found, this seems the best site for CO2 sampling.
An interesting alternative, even if slightly less accurate, is represented by yCO2, i.e. by reading the etCO2 at the Y-piece. Of note, the value of etCO2 at the Y-piece is much lower than the actual etCO2 of the patient, because the expired CO2 dilutes within the helmet internal volume, which is much greater than the patient's expiratory volume. Therefore, yCO2 can be used to estimate CO2 rebreathing but not to estimate the patient's arterial CO2.
Usually CO2 rebreathing had been evaluated by means of the end-inspiratory value at the airway opening [1, 13, 15, 16], but in our experience eiCO2 grossly underestimated mean inspiratory CO2, particularly when it was significant. The reason was the slow decrease in inhaled CO2 concentration during the inspiratory phase of helmet ventilation.
When a steady state is reached and CO2 is stable inside the helmet, the amount of CO2 entering the helmet per minute must be equal to the amount of CO2 leaving the device in the same time. Assuming a non-significant dead space at the Y-piece of the ventilator circuit, the only source of CO2 entering the helmet is CO2 production by the patient. Concerning the CO2 leaving the system, that is a function of the mean CO2 inside the helmet and the sum of all flows directed from inside the helmet to outside: namely the expired (by patient and helmet) ventilation through the Y-piece, plus the air leaks at the collar–neck interface, plus, if present, the outflow of an additional flow-by. Accordingly, the theoretical equation of CO2 steady state inside the helmet is:
$$
\begin{aligned}
{\text{V}}^{\prime}{\text{CO}}_2 &={\text{hCO}}_2\times {\text{MV}}_{\text{total}}, {\text{or }}\\ {\text{ hCO}}_2 &={\text{V}}^{\prime}{\text{CO}}_2 / {\text{MV}}_{\text{total}}.
\end{aligned}
$$
Therefore, the mean CO2 concentration inside the helmet should be directly related to V'CO2, inversely related to MVtotal and unaffected by any other factor. This hypothesis was fully confirmed by our experimental data (Fig. 4 and 5): of note, during phasic helmet ventilation we obtained results similar to those observed by Taccone et al. during continuous-flow CPAP [13].
As a matter of fact, all the manipulations tested in our second series affected CO2 rebreathing insofar as they were able to change the total flow passing through the helmet.
The increase of inspiratory pressure above PEEP produced a progressive increase in MVtotal; accordingly, hCO2 progressively decreased. In contrast with our findings, studies on healthy volunteers recently found no relationship between pressure support level and CO2 rebreathing [15, 16]. A possible explanation is that, in order to keep their minute ventilation constant, volunteers may react to an increase in pressure support with a decrease in spontaneous inspiratory activity, thus limiting the effect on MVtotal. Most importantly, the theoretical relationship between MVtotal and hCO2 (Fig. 4) tends to flatten for high values of MVtotal, and the flattening takes place for lower values of MVtotal when V'CO2 is lower. In the study by Costa and coworkers [15], the mean MVtotal was close to the top of our second series, while V'CO2 was much lower. Therefore, in these conditions the expected variations in helmet CO2 due to changes in MVtotal were really low. Finally, rebreathing was assessed by the end-inspiratory CO2 at the airway opening in these studies [15, 16], i.e. by a parameter we have found to have major limitations.
The relative distribution of MVtotal between patient, helmet and air leaks does not affect rebreathing. In confirmation of that, when we selected a low compliance in the lung model, this resulted in a large change in the patient:helmet ventilation ratio (from 2:1 to about 1:1) with only a small change in hCO2, the latter fully explained by a proportionate change in MVtotal.
With regard to leaks, these can decrease CO2 inside the helmet by an increase in MVtotal. In any case, a helmet provided with an “intentional” leakage port will decrease helmet pressurization and patient's inspiratory assistance, probably turning out counterproductive for CO2 removal.
Concerning the role of helmet volume, inflation of the cushions resulted in a 33% volume reduction but in our experience was associated with no change in hCO2. In clinical practice, the inflated cushions may facilitate the initial pressurization of the helmet, improve the comfort of the patient and stabilize the system, eventually limiting air leaks: the overall effect on rebreathing is difficult to predict, but probably of low significance.
The addition of a flow-by through the helmet was very effective in clearing CO2 around the head of the patient. However, the flow-by system we used in our experimental setting is not very practical for clinical application. An interesting option might be the use of the bias flow of the ventilator as a flow-by. In order to force the bias flow to pass through the helmet, the ventilator must be connected to the helmet by two independent ports, for inspiration and expiration. Moreover, for a significant effect on CO2 rebreathing the bias flow should be continuous or at least applied during the entire expiratory phase, and it should be adjustable at relatively high values, at least 10 l/min.
This study has some limitations. We performed a bench study, because the pure effects of changes of CO2 production or minute ventilation on rebreathing would have been difficult to study in patients or volunteers. The lung model was passive and the ventilation controlled, in contrast to clinical practice. The use of an active physical model would have offered the opportunity, through modifications of the spontaneous activity, to change both the patient ventilation and the patient:helmet ventilation ratio. In our experimental setting, we obtained a similar effect by changing the lung model compliance. Finally, just one type of helmet was studied. Helmet models can differ in elastic properties, tightness or inner volume: these factors are related to rebreathing through their influence on helmet ventilation and leaks or, in the case of volume, not related at all.