The most important functional characteristic of a helmet in the context of concussion is impact energy attenuation; a characteristic that has also been referred to as acceleration management. Ideally, the impact will cause the helmet to deform a substantial proportion of its thickness, without fully deforming or “bottoming out.” The liner of the helmet or, in the case of padded headgear worn in rugby union the entire helmet, largely determines the impact energy attenuation performance. In short, the greater the deformation of the helmet the greater the reduction in impact force as well as in head acceleration. The helmet can also distribute the impact force over an area larger than the contact area. In helmets with a well-established role in transport and sport, e.g., bicycle and motorcycle helmets, the helmet is designed for a single crash event. In contrast, American football, rugby union, and ice hockey helmets are designed to provide protection throughout a season or more of multiple head impact exposures. The general properties of helmets and their function have been addressed well by many authors, e.g., Newman (1993) and Hoshizaki and Brien (2004), and will not be repeated in this chapter [11, 12].

The next most important functional characteristics are the mass, mass distribution, fit, restraint system, and vision. Sports helmets need to be wearable during extreme physical activities; therefore, helmet mass must be minimized. The mass distribution of the helmet and attachments is important in reducing the flexion moment that the helmet may apply to the head and neck. A flexion moment will be counteracted by neck extensor activation leading to muscle fatigue and increased joint reaction forces. It is imperative to ensure that the helmet and all components are correctly selected and adjusted for the individual athlete. Providing a kit bag with a few helmets to fit all the team is not best practice. Vision and the restraint system characteristics are usually addressed in sports helmet standards. Where faceguards (visors) are mounted to helmets to prevent projectile to face or head impacts, the adjustment of the faceguard is important as apertures may permit a projectile travelling at speed to strike the face directly. Some helmet styles permit adjustment of the faceguard, e.g., cricket helmets, where others do not, e.g., baseball helmets. The former introduces the potential for injury due to misuse.

Helmet performance is assessed in the laboratory by examining the capacity of the helmet to minimize headform acceleration in impact tests. These tests are conducted against set criteria, e.g., a linear acceleration pass criterion, or to derive an injury risk estimate. During a test a selected amount of impact energy is delivered to the helmet–headform system via a drop rig, pendulum, or mechanical device. The headform’s linear and, in some cases, angular acceleration is measured during the impact. The input characteristics of the tests, e.g., energy, dimensions of impact interface, and headform, have gradually evolved to reflect knowledge on impact exposures in specific sports. The output characteristics, e.g., headform dynamic responses, have also evolved to reflect knowledge on injury mechanisms and human tolerance. However, requirements in many helmet standards are not currently aligned to maximize the potential for standard compliant helmets to prevent concussion. This would require the lowering of pass levels, e.g., to well below 100 g, and consideration for angular acceleration criteria and related test methods. As will be presented in this section, more valid assessments of helmet performance are observed when the laboratory tests reflect the impact exposures in the specific sport (impact location, impact severity, interface characteristics, and frequency) and the biofidelity of the head–neck system is considered. A range of headforms is used in research and standards testing: Hybrid III headforms, rigid ISO headforms, and NOCSAE headforms. Each has a distinct influence on the test outcomes. In an otherwise equivalent impact, head acceleration will be greater with a rigid ISO headform in comparison to a Hybrid III headform.

The author and others have conducted baseline tests on bare headforms. These reveal a clear risk of concussion related to linear head acceleration even in impacts equivalent to the head falling 0.5 m [3]:

In the context of laboratory impact tests, helmets need to reduce both linear and angular headform acceleration. As a guide the 15 % likelihood of concussion for adult males is 45 g and the 50 % likelihood is 75 g for resultant linear acceleration at the head’s center of gravity [16]. For the bare headform impacts described above helmets need to reduce the linear acceleration in the range two- to tenfold to prevent concussion. Angular acceleration tolerance thresholds vary; Rowson et al. reported in 2012 that the 75 % likelihood of concussion for resultant angular acceleration is 6.9 krad/s² [17].

Rugby—Helmets (padded headgear) in rugby must comply with the International Rugby Board’s (IRB) performance regulations [18]. The helmet properties are restricted to an undeformed thickness of 10 mm and a foam density of 45 kg/m³. The IRB’s impact performance requirements state that in a 13.8 J rigid (EN 960) headform impact onto a rigid flat anvil the peak headform acceleration shall not be less than 200 g. The mandated performance requirements exclude headgear from preventing concussion due to the biomechanical criteria and are inconsistent with the philosophy of many helmet standards.

Impact tests on helmets meeting the IRB’s requirements (“standard”) and a “modified” version were conducted by the author [19]. The modified headgear was 16 mm thick and made from 60 kg/m³ polyethylene foam. The standard headgear was 10 mm thick and made from 45 kg/m³ polyethylene foam. Tests using a rigid headform from a 0.3 m drop height produced peak accelerations in the range 276–689 g for standard headgear and 69–123 for modified headgear. At 0.4 m peak accelerations for the modified headgear were 110–273 g. Figure 9.1 shows a level of consistency between the laboratory tests and the results of the randomized controlled trial superimposed onto injury likelihood curves [16]. The performance of the modified headgear in laboratory tests identified a potential in low severity impacts for the headgear to reduce the linear acceleration to a tolerable range. In the epidemiological study there was a greater than 50 % non-significant reduction in missed game concussions based on a compliance analysis. With greater compliance, this may have been a significant association. The epidemiological study was of players aged under 13 years to under 20 years; therefore, the tolerance data presented in Fig. 9.1 may not be suitable, but the impact exposure may be less severe than observed in adult rugby union and Australian football and more reflective of the laboratory tests, noting also that a rigid headform was used. These highlight the intrinsic and extrinsic factors described in the introduction.