3. VRU accidentology and biomechanics

Pedestrian and cyclist protection engineering begins with a deceptively simple question: what actually happens to a human body when a vehicle strikes it? Answering that question requires three connected disciplines. Accidentology describes the kinematics of the collision — how the body is loaded, at what speed, and against which vehicle structures. Biomechanics translates those loads into tissue-level injury mechanisms. Injury metrics convert physical measurements into scalar values that correlate with real-world injury risk, so that a vehicle can be designed, tested, and regulated against a numeric threshold.

This chapter builds that chain from the ground up, using the head and the lower extremity as the two worked examples, because they dominate the VRU injury burden and because their metrics behave very differently.

Collision Kinematics: The Loading Sequence

For a typical passenger car front end, a pedestrian impact at 40 km/h unfolds in roughly 100–200 ms. The sequence is highly structured and largely determined by vehicle geometry rather than by the pedestrian's reaction, which is why it can be standardized in regulation.

• Bumper contact (0–20 ms): the bumper beam and fascia load the tibia and knee region. Bumper height and stiffness dominate the outcome here.
• Thigh–bonnet leading edge contact (20–60 ms): the pelvis or upper leg rotates about the contact point, and the torso begins to rotate toward the bonnet.
• Head–bonnet or head–windscreen contact (80–180 ms): the head strikes the bonnet top, A-pillar, or windscreen, depending on vehicle front-end shape and impact speed.
• Secondary ground contact: after the primary impact, the body is thrown and strikes the road, adding a second injury phase that is not covered by the vehicle test but is significant in real accidents.

Two consequences follow directly. First, the wrap-around distance (WAD) — the arc length from the ground reference line to a point on the bonnet — determines which body region reaches which vehicle structure. Second, the bonnet leading edge height controls whether the impact is primarily a leg-bending or a pelvis-loading event. These are geometric design parameters, not crash-test artifacts.

Injury Mechanisms at Tissue Level

Injury is not caused by force alone; it is caused by the tissue-level response to force. Four mechanisms cover most VRU injuries:

• Bone fracture under bending and compression. The tibia and femur fail when bending moment or axial compression exceeds the bone's tolerance. Tibia fracture is the classic bumper-induced injury.
• Ligament rupture under shear and bending. The knee is vulnerable to combined valgus/varus bending plus shear, which ruptures the collateral ligaments and, in severe cases, the cruciate ligaments. This is the mechanism behind the knee-bending angle criterion.
• Brain injury from linear and rotational acceleration. Linear acceleration produces focal contusion and skull fracture; rotational acceleration produces diffuse axonal injury (DAI) and subdural hematoma. This distinction is why HIC alone is insufficient and why rotational criteria such as BrIC have been introduced.
• Soft-tissue and vascular injury. Aortic rupture, liver laceration, and pelvic fracture occur in higher-severity impacts, often from the secondary ground contact.

AIS: The Common Language of Injury Severity

The Abbreviated Injury Scale (AIS) is an anatomically based, consensus-driven ordinal scale from 1 (minor) to 6 (currently untreatable). It is not a linear scale, and it is not a probability. Its value is that it gives engineers and clinicians a shared vocabulary, and it allows the Maximum AIS (MAIS) and the Injury Severity Score (ISS) to be computed for a casualty.

In regulation, the target is usually expressed as a probability of exceeding a given AIS level. For example, a head criterion threshold is often set to correspond to a 50th-percentile risk of AIS 3+ head injury, while lower-leg criteria are set against AIS 2+ tibia fracture risk. The metric is the proxy; the AIS level is the actual safety objective.

HIC: Construction and Interpretation

The Head Injury Criterion (HIC) is defined as the maximum over all time windows of the expression:

HIC = max[ (t2 - t1) * ( (1/(t2-t1)) * ∫ a(t) dt )^2.5 ]

where:
  a(t) = resultant translational acceleration at the head CG,
         expressed in g (1 g = 9.81 m/s²)
  t1, t2 = window boundaries, typically limited to 15 ms
           (HIC15) or 36 ms (HIC36)

Three practical points matter when using HIC:

1. It is a translational criterion. HIC is computed from the resultant linear acceleration only. It is blind to rotation, which is why a pure HIC-based assessment can pass a case that produces DAI.
2. The window length changes the number. HIC15 and HIC36 are not interchangeable. A short, high-amplitude pulse can produce a high HIC15 but a lower HIC36. Always state the window.
3. It is a risk correlate, not a physical injury. A HIC of 1000 is conventionally associated with roughly a 15–20% risk of AIS 3+ head injury for a mid-sized adult, but this mapping is population-based and carries substantial uncertainty.

Lower Extremity Metrics

Lower extremity criteria are expressed as physical quantities measured on the legform impactor, not as a single composite index. The three primary quantities are:

• Tibia acceleration (measured at the upper and lower tibia), in g. This correlates with tibia fracture.
• Knee bending angle, in degrees, measured between the femur and tibia axes. This correlates with ligament rupture.
• Knee shear displacement, in millimeters, measured between the femur and tibia. This correlates with combined ligament and joint injury.

A representative regulatory-style acceptance set for a legform test is shown below. These values are illustrative of the structure of such criteria and should be confirmed against the applicable regulation before use.

The engineering logic is straightforward: reduce bumper stiffness and tune bumper height so that the legform's tibia acceleration and knee bending stay below these limits. A stiffer bumper raises tibia acceleration; a bumper positioned too high or too low shifts the load path and increases knee bending.

Worked Example: Reading a Legform Test Result

Consider a legform test at 40 km/h against a vehicle front end. The measured channels are:

Upper tibia acceleration peak : 158 g
Lower tibia acceleration peak : 212 g
Knee bending angle peak       : 14.2°
Knee shear displacement peak  : 4.1 mm

The upper tibia, bending angle, and shear channels pass. The lower tibia channel fails at 212 g against a 200 g limit. The correct diagnostic step is to inspect the lower bumper geometry: a stiff lower stiffener or a low bumper beam is the usual cause, because it concentrates load on the lower tibia. The typical remedy is to soften or reposition the lower leg support, not to change the upper bumper. Re-running the test after the change should show the lower tibia peak dropping while the other three channels remain within limits.

From Metrics to Design Decisions

The metrics are only useful if they drive geometry. The figure below summarizes how the four legform channels and the head criterion map back to specific vehicle front-end parameters, and how a failing channel points to a specific design change.

The mapping is not one-to-one in practice — changing bumper beam height affects both upper and lower tibia acceleration — but the figure captures the diagnostic logic: identify the failing channel, trace it to the dominant geometric parameter, change that parameter, and re-evaluate all channels. This is the working method that connects accidentology and biomechanics to a design decision.