Safety helmets are currently designed according to standards that aim at reducing the maximum acceleration transmitted to the centre of gravity of the head.
Figure 1(a) below shows the main components of a motorcycle safety helmet: (1) shell, made of plastics or composite materials, (2) energy absorbing liner, usually made of polystyrene in motorcycle helmets, (3) comfort padding and (4) retention system.
The standards currently adopted in Europe assess the protective performance of a helmet by imposing to it a series of impacts, as shown in figures 1(b) and 1(c) for the case of motorcycle helmets. A metal head-form, shown in red in figure 1(c), wearing the helmet, is raised to a certain height and then it falls on an anvil with a given speed. Every motorcycle helmet is subjected to four impacts: on the front (as shown in figure 1(c)), the top, the back and the side and in all cases the peak linear acceleration at the centre of mass of the head-form has to be below a given limit. At the same time also the duration of the impact is measured so that high acceleration values can affect the head-form only for limited durations.
The loads applied to the head are usually defined in terms of linear and rotational accelerations: the linear acceleration is mainly generated by the force component normal to the impact surface and the rotational acceleration by the tangent component, as shown in figure 2. A considerable amount of research has been devoted to describe how the characteristics of the linear and rotational acceleration influence specific traumatic brain injury lesions [7].
Figure 2: if the impact direction is perpendicular to the impact surface (top) the main acceleration component is ‘linear’, if instead the impact direction has a tangential component (bottom) then there is also an appreciable rotational acceleration [Svein Kleiven, ‘Why most traumatic brain injuries are not caused by linear acceleration but skull fractures are’, Front. Bioeng. Biotechnol., 07 November 2013 | https://doi.org/10.3389/fbioe.2013.00015.].
The mechanical response of the brain is strictly linked to the kinematics of the event that caused the head impact and the consequent damage will depend on the final result of the complex interrelation between linear and angular acceleration strains, ultimately acting on the internal structures of the brain. It is then essential to consider that the linear and rotational acceleration components will determine specific traumatic brain injury lesions and to examine the influence of the characteristics of such pulses and how they account for the variance in predicting the outcome of the main TBI lesions, namely: contusion, subdural hematoma (SDH), subarachnoid haemorrhage (SAH), and epidural hematoma (EDH). Recent studies pointed out for example that the best predictors for brain contusions seem to be “time to peak” resultant and component in the x and y characteristics of the linear and angular acceleration. The total duration of the pulse is also important while the influence of the z axis (see figure 2) on the variance was fond to be low probably because on this data set there were no “top of the head” impacts [7].
When considering SDH, it has been stressed that time based variables such as time to peak for the linear and angular acceleration in the x and y components accounted for a large part of the variance. Total duration was also found to be influential. Interestingly, in SDH, more variance was accounted for by the angular components on the x, y, and z axes (30%). These data revealed that, in generating a subdural hematoma, the angular acceleration characteristics accounted for more variance than the linear acceleration characteristics. This finding is strongly in agreement with previous literature which has identified subdural hematoma as a rotationally-influenced lesion as opposed to a linear dominant injury similar to other forms of TBI [8]. Post-traumatic subarachnoid haemorrhage is more likely to be accounted for by time to peak and total duration characteristics. The remaining components involve the x and z linear curve characteristics and the x, y, and z angular curve characteristics. Like subdural hematoma, a large amount of the variance, after the time-based characteristics, is accounted for by angular loading curve peak, slope, and integral. This indicates a rotational influence for the risk of subarachnoid haemorrhage which is based within the x, y, and z components and not the characteristics of the resultant linear or angular acceleration.
On the contrary, Epidural Hematomas are the consequence mainly of just two components which accounted for 94.6% of the total variance. The first component is time based, and similar to the other TBI lesions. The second component is mostly the x and y axes linear dynamic response characteristics. It is then important to understand that in order to predict the gravity of an impact in terms of brain lesions is of fundamental importance to analyze the singular components of the linear and angular acceleration loading curve on the x, y, and z axes, rather than the resultant values alone. A large body of evidence is showing that severe lesions are due to extremely short duration time to peak and duration of linear and angular accelerations. The intervention of protective devices (helmets, crash pads etc.) serve to lengthen this response, but also to prevent adverse kinematics of the head and brain.
It is apparent that the set-up for helmet standard tests is reasonable but it does not take into account several aspects of real life accidents and their wide variety. The main way, which is currently followed, to improve the protective capability of the helmets is to reduce the threshold values of the quantities measured in standard impacts, such as the peak linear acceleration. Standard tests do not account for the presence of the body, do not take into account rotational accelerations and, most of all, do not consider the dynamics of human organs and tissues in the skull. All these limitations are mainly due to the fact that the currently adopted standards were discussed and came into force several decades ago and they reflect the technological capabilities of that time. There are proposals to include limitations to the maximum value of the angular acceleration in the standards [9], but it is clear that global quantities such as peak linear or angular accelerations, impact duration or time to maximum acceleration are not sufficient to define an impact. Reference [7] stresses that dissimilar acceleration loading curves with identical peak values and time integral produced different magnitudes of brain deformation and therefore are likely to induce different injuries. Therefore, to reduce the consequences of head impacts and prevent brain trauma, understanding the mechanism of injury is essential. The incidence of brain injury is linked to how the kinematics of a brain injury event affects the internal structures of the brain. A complete approach for predicting the incidence of TBI in humans requires several steps [1]:
(1) defining the external mechanical loads experienced by the head during situations that cause injury,
(2) using models of the brain to estimate how these external mechanical loads transfer to mechanical conditions (e.g., strain, strain rate, stress, etc.) in the brain at the tissue scale, and
(3) using tissue tolerance criteria to determine the regions of the brain that will be injured or impaired as a result of the external applied loading.
REDIPhE is a research project by
REDIPhE project is funded by
