## 4 Discussion
This finding could be ascribed to the lower duration of the swing phase of gait initiation under the FFS compared to the RFS condition (\(\Delta=6.70\%\)), which implies that the period during which the centre of mass was falling laterally under gravity was reduced. In turn, this shorter duration of the swing phase under the FFS condition could be ascribed to the lower force braking under the FFS condition (\(\Delta=19.4\%\)), which accelerated the fall of the centre of mass to a greater extent (as discussed above). As a second factor contributing to enhanced stability, the size of the ML base of support was increased under the FFS compared to the RFS condition. This result could not be ascribed to a strategy of a more lateral placement of the swing foot [24; 35] (since the distance between the two heels remained unchanged between the two foot strike conditions), but to foot anatomy. Under the FFS condition, the ML base of support size corresponded to the distance between the stance heel and the fifth metatarsal of the swing foot, while under the RFS condition, it corresponded to the distance between the two heels. It was therefore smaller than under the FFS condition. This difference reached up to 10 cm in some subjects. The question arises as to why the amplitude of ML APAs is decreased when stability is increased with the change in the foot strike pattern. Generating ML APAs is known to have an energetic cost, requiring a coordinated activation of hip abductors and ankle dorsiflexors [7; 42]. Although not shown in the literature to date, it may also have an attentional cost, as has previously been documented for the control of static erect posture and level walking [43; 44]. An optimal strategy for stability control would appear to be to reduce this cost by attenuating the ML APAs under the FFS condition while, at the same time, increasing stability with the mechanical and anatomical changes reported above--neither of which requires any additional energetic or attentional cost (even less for attenuated force braking). This optimal strategy implies that one (or both) of these factors associated with the foot strike pattern be accounted for in the programming of ML APAs. This statement further reinforces the hypothesis proposed above of an interdependent relationship between motor mechanisms for optimal stability control. Interestingly, previous studies focusing on the effect of different sorts of constraints applied to the postural system during gait initiation (such as temporal pressure [41; 45], progression velocity [24], the presence of an obstacle of different heights and distances to be cleared [10; 11], the addition of a mass distributed (a)symmetrically over the body or the ankles [46; 47; 48], etc.) repeatedly showed that the CNS modulated the spatiotemporal features of ML APAs so as to maintain an equivalent MOS value. This motor invariance through these disparate postural constraints led the authors to suggest that the CNS would set a fixed MOS value before initiating gait and would then adjust ML APAs features to keep it unchanged (for review, see [2]). Thus, the present finding that the MOS was not invariant is not in line with these previous results and related hypothesis. Now, it may be recalled that the participants systematically used an RFS pattern in all of the previous studies. Therefore, the MOS was computed and compared across conditions at the moment the swing heel hit the ground. By contrast, the present study compared the MOS of two different postures of the landing foot. These foot postures were responsible for the difference in the ML size of the base of support, which was a major factor explaining the MOS differences between the two foot strike conditions. It is possible that, for a given foot strike pattern (RFS or FFS), a fixed MOS value is planned before initiating gait, and may be greater for the FFS than the RFS pattern due to foot anatomy.