2 Normal gait
In order to understand pathological gait, it is necessary first to understand normal gait, since this provides the standard against which the gait of a patient can be judged. However, there are two pitfalls which need to be borne in mind when using this approach. Firstly, the term ‘normal’ covers both sexes, a wide range of ages and an even wider range of extremes of body geometry, so that an appropriate ‘normal’ standard needs to be chosen for the individual who is being studied. If results from an elderly female patient are compared with normal data obtained from physically fit young men, there will undoubtedly be large differences, whereas comparison with data from healthy elderly women may show the patient’s gait to be well within normal limits which are appropriate to her sex and age. The second pitfall is that even though a patient’s gait differs in some way from normal, it does not follow that this is in any way undesirable or that efforts should be made to turn it into a ‘normal’ gait. Many gait abnormalities are a compensation for some problem experienced by the patient and, although abnormal, they are nonetheless useful.
Having said all that, it is very important to understand normal gait and the terminology which is used to describe it, before going on to look at pathological gait. The chapter starts with a very brief historical review and then gives an overview of the gait cycle, before going on to study in detail how the different parts of the locomotor system are used in walking.
As walking is such a familiar activity, it is difficult to define it without sounding pompous. However, it would be remiss not to attempt a definition. Normal human walking and running can be defined as ‘a method of locomotion involving the use of the two legs, alternately, to provide both support and propulsion’. In order to exclude running, we must add ‘at least one foot being in contact with the ground at all times’. Unfortunately, this definition excludes some forms of pathological gait which are generally regarded as being forms of walking, such as the ‘three-point step-through gait’ (see Fig. 3.21), in which there is an alternate use of two crutches and either one or two legs. It is probably both unreasonable and pointless to attempt a definition of walking which will apply to all cases – at least in a single sentence!
Gait is no easier to define than walking, many dictionaries regarding it as a word primarily for use in connection with horses! This is understandable, since quadruped animals have a repertoire of natural gaits (walking, trotting, cantering, galloping, etc.), as well as some artificial ones, such as that learned by ‘Tennessee Walking Horses’ in the area where one of the authors lives. Most people tend to use the words gait and walking interchangeably. However, there is a difference: the word gait describes ‘the manner or style of walking’, rather than the walking process itself. It thus makes more sense to talk about a difference in gait between two individuals than about a difference in walking.
The history of gait analysis has shown a steady progression from early descriptive studies, through increasingly sophisticated methods of measurement, to mathematical analysis and mathematical modelling. Only a brief account of the development of the discipline will be given here. Good reviews of the early years of gait analysis have been given by Garrison (1929), Bresler and Frankel (1950) and Steindler (1953). The more recent history of gait analysis, and of clinical gait analysis in particular, was covered in three excellent review papers by Sutherland (2001, 2002, 2005).
Walking has undoubtedly been observed ever since the time of the first men, but the systematic study of gait appears to date from the Renaissance when Leonardo da Vinci, Galileo and Newton all gave useful descriptions of walking. The earliest account using a truly scientific approach was in the classic De Motu Animalum, published in 1682 by Borelli, who worked in Italy and was a student of Galileo. Borelli measured the centre of gravity of the body and described how balance is maintained in walking by constant forward movement of the supporting area provided by the feet. The Weber brothers in Germany gave the first clear description of the gait cycle in 1836. They made accurate measurements of the timing of gait and of the pendulum-like swinging of the leg of a cadaver.
Two pioneers of kinematic measurement worked on opposite sides of the Atlantic in the 1870s. Marey, working in Paris, published a study of human limb movements in 1873. He made multiple photographic exposures, on a single plate, of a subject who was dressed in black, except for brightly illuminated stripes on the limbs. He also investigated the path of the centre of gravity of the body and the pressure beneath the foot. Eadweard Muybridge (born in England as Edward Muggeridge) became famous in California in 1878 by demonstrating that, when a horse is trotting, there are times when it has all four of its feet off the ground at once. The measurements were made using 24 cameras, triggered in quick succession as the horse ran into thin wires stretched across the track. In the next few years, Muybridge made a further series of studies, of naked human beings walking, running and performing a surprising variety of other activities!
The most serious application of the science of mechanics to human gait during the nineteenth century was the publication in Germany, in 1895, of Der Gang des Menschen, by Braune and Fischer. They employed a technique similar to Marey’s, but using fluorescent strip-lights on the limbs instead of white stripes. The resulting photographs were used to determine the three-dimensional trajectories, velocities and accelerations of the body segments. Knowing the masses and accelerations of the body segments, they were then able to estimate the forces involved at all stages during the walking cycle.
Further valuable work on the dynamics of locomotion was done by Bernstein in Moscow in the 1930s. He developed a variety of photographic techniques for kinematic measurement and studied over 150 subjects. Particular attention was paid to the centre of gravity of the individual limb segments and of the body as a whole.
Further progress followed the development of the force platform (also called the forceplate). This instrument has contributed greatly to the scientific study of gait and is now standard equipment in gait laboratories. It measures the direction and magnitude of the ground reaction force beneath the foot. An early design was described by Amar in 1924 and an improved one by Elftman in 1938. Both were purely mechanical, the force applied to the platform causing the movement of a pointer. In Elftman’s design the pointers were photographed by a high-speed movie camera.
For a full understanding of normal gait, it is necessary to know which muscles are active during the different parts of the gait cycle. The role of the muscles was studied by Scherb, in Switzerland, during the 1940s, initially by palpating the muscles as his subject walked on a treadmill, then later by the use of electromyography (EMG). Further advances in the understanding of muscle activity and many other aspects of normal gait were made during the 1940s and 1950s by a very active group working in the University of California at San Francisco and Berkeley, notable among whom was Verne Inman. This group later went on to write Human Walking (Inman et al., 1981), published just after Inman’s death, which to many people is the definitive textbook on normal gait. This has now gone through several editions, the latest edition of this being Rose and Gamble (2005). Another classic text on EMG is Muscles Alive: Their Functions Revealed by Electromyography by John Basmajian, which unfortunately has not been updated since 1985.
The use of EMG in gait analysis has received much attention but perhaps the most influential paper published was ‘The use of surface electromyography in biomechanics’ by Carlo De Luca in 1997, which gave a summary of recommendations but perhaps more importantly a summary of problems which at the time needed resolution. Further standardisation was achieved through the SENIAM (Surface ElectroMyoGraphy for the Non-Invasive Assessment of Muscles) project coordinated and managed by Hermie Hermens and Bart Freriks from Enschede, which is now considered by many as the definitive recommendations for electrode configuration and sensor positioning on specific muscles.
A major contribution to the mechanical analysis of walking, also from the Californian group, was made by Bresler and Frankel (1950). They performed free-body calculations for the hip, knee and ankle joints, allowing for ground reaction forces, the effects of gravity on the limb segments and the inertial forces. The analytical techniques developed by these workers formed the basis of many current methods of modelling and analysis.
An important paper describing the possible mechanisms which the body uses to minimise energy consumption in walking, again from California, was published by Saunders et al. (1953). Further important work on energy consumption and in particular the energy transfers between the body segments in walking, was published by Cavagna and Margaria (1966), working in Italy. By 1960, research began to concentrate on the variability of walking, the development of gait in children and the deterioration of gait in old age. Patricia Murray, working in Milwaukee, Wisconsin, published a series of papers on these subjects, including a detailed review (Murray, 1967).
Once the motions of the body segments and the actions of the different muscles had been examined and documented, attention passed to the forces generated across the joints. Although limited calculations of this type had been made previously, the study by Paul (1965) was the first detailed analysis of hip joint forces during walking. A subsequent paper by the same author also included an analysis of the forces in the knee (Paul, 1966). Since then there have been many mathematical studies of force generation and transmission across the hip, knee and ankle.
The 1970s and 1980s saw great improvements in methods of measurement. The development of more convenient kinematic systems, based on electronics rather than photography, meant that results could be produced in minutes rather than days. Reliable force platforms with a high-frequency response became available, as well as convenient and reliable EMG systems. The availability of high-quality three-dimensional data on the kinetics and kinematics of walking, and the ease of access to powerful computers, made it possible to develop increasingly sophisticated mathematical models. Gait laboratories now routinely measure joint moments and powers for the hip, knee and ankle. Somewhat less reliably, estimates can also be made of muscle, ligament and joint contact forces.
The 1990s and 2000s have seen the emergence of increasingly powerful systems, higher speed cameras, greater portability, smaller markers and a wide variety of marker sets, and larger volumes of area that data can be attained from.
From the earliest days, it has been the hope of most of those working in this field that gait measurements would be found useful in the management of patients with walking disorders. Many of the early workers made studies of people who walked with abnormal gait patterns and some (notably Amar, Scherb and the Californian group) attempted to use the results for the benefit of individual patients. However, the results were not particularly impressive.
Since 1960, there has been a more serious attempt to take gait analysis out of the research laboratory and into the clinic. With the improvements in measurement and analytical techniques, the major limitation now is not the ability to produce high-quality data but knowing how best to use these data for the benefit of patients. It is fair to say that in the early days, far more progress was made in scientific gait analysis, particularly as applied to normal subjects, than in the application of these techniques for the benefit of those with gait disorders. However, since about 1980, there has been a steady increase in the effective use of gait analysis in the clinical management of patients.
As well as a gradual increase in the clinical use of scientific gait analysis, there has also been a growing interest in the use of observational or visual gait analysis. This has become much easier to perform since digital video cameras have become widely available, which is reviewed in more detail by Sutherland (2001, 2002, 2005).
The gait cycle is defined as the time interval between two successive occurrences of one of the repetitive events of walking. Although any event could be chosen to define the gait cycle, it is generally convenient to use the instant at which one foot contacts the ground (‘initial contact’). If it is decided to start with initial contact of the right foot, as shown in Figure 2.1, then the cycle will continue until the right foot contacts the ground again. The left foot, of course, goes through exactly the same series of events as the right, but displaced in time by half a cycle.
These seven events subdivide the gait cycle into seven periods, four of which occur in the stance phase, when the foot is on the ground, and three in the swing phase, when the foot is moving forward through the air (Fig. 2.1). The stance phase, which is also called the ‘support phase’ or ‘contact phase’, lasts from initial contact to toe off. It is subdivided into:
The duration of a complete gait cycle is known as the cycle time, which is divided into stance time and swing time. Unfortunately, the nomenclature used to describe the gait cycle varies considerably from one publication to another. The present text attempts to use terms which will be understood by most people working in the field; alternative terminology will be given where appropriate. Wall et al. (1987) pointed out that the usual terminology is inadequate to describe some severely pathological gaits.
Figure 2.2 shows the timings of initial contact and toe off for both feet during a little more than one gait cycle. Right initial contact occurs while the left foot is still on the ground and there is a period of double support (also known as ‘double limb stance’) between initial contact on the right and toe off on the left. During the swing phase on the left side, only the right foot is on the ground, giving a period of right single support (or ‘single limb stance’), which ends with initial contact by the left foot. There is then another period of double support, until toe off on the right side. Left single support corresponds to the right swing phase and the cycle ends with the next initial contact on the right.
In each double support phase, one foot is forward, having just landed on the ground, and the other one is backward, being just about to leave the ground. When it is necessary to distinguish between the two legs in the double support phase, the leg in front is usually known as the ‘leading’ leg and the leg behind as the ‘trailing’ leg. The leading leg is in ‘loading response’, sometimes referred to as ‘braking double support’, ‘initial double support’ or ‘weight acceptance’. The trailing leg is in ‘pre-swing’, also known as ‘second’, ‘terminal’ or ‘thrusting’ double support or ‘weight release’.
In each gait cycle, there are thus two periods of double support and two periods of single support. The stance phase usually lasts about 60% of the cycle, the swing phase about 40% and each period of double support about 10%. However, this varies with the speed of walking, the swing phase becoming proportionately longer and the stance phase and double support phases shorter, as the speed increases (Murray, 1967). The final disappearance of the double support phase marks the transition from walking to running. Between successive steps in running there is a flight phase, also known as the ‘float’, ‘double-float’ or ‘non-support’ phase, when neither foot is on the ground. A detailed study of gait cycle timing was published by Blanc et al. (1999).
The terms used to describe the placement of the feet on the ground are shown in Figure 2.3. The stride length is the distance between two successive placements of the same foot. It consists of two step lengths, left and right, each of which is the distance by which the named foot moves forward in front of the other one. In pathological gait, it is common for the two step lengths to be different. If the left foot is moved forward to take a step and the right one is brought up beside it, rather than in front of it, the right step length will be zero. It is even possible for the step length on one side to be negative, for example, if the left foot never catches up with the right foot, the distance between the left and right feet will be negative. However, the stride length starting with the left heel strike must always be the same as the stride length starting with the right heelstrike, unless the subject is walking around a curve where the inside leg will have a shorter stride length than the outside leg. This definition of a ‘stride’, consisting of one ‘step’ by each foot, breaks down in some pathological gaits, in which one foot makes a series of ‘hopping’ movements while the other is in the air (Wall et al., 1987). There is no satisfactory nomenclature to deal with this situation.
The walking base (also known as the ‘stride width’ or ‘base of support’) is the side-to-side distance between the line of the two feet, usually measured at the midpoint of the back of the heel but sometimes below the centre of the ankle joint. The preferred unit for stride length and step length is the metre and for the walking base, millimetres. The pattern of walking known as ‘tandem gait’ involves walking with the heel of one foot placed directly in front of the toes of the other, i.e. with a walking base close to zero. Although this pattern is not typically seen, even as a pathological gait, it requires good balance and coordination and it is often used by the police as a test for intoxication!
The toe out (or, less commonly, toe in) is the angle in degrees between the direction of progression and a reference line on the sole of the foot. The reference line varies from one study to another; it may be defined anatomically but is commonly the midline of the foot, as judged by eye.
It is common experience that you need to walk more carefully on ice than on asphalt. Whether or not the foot slips during walking depends on two things: the coefficient of friction between the foot and the ground, and the relationship between the vertical force and the forces parallel to the walking surface (front-to-back and side-to-side). The ratio of the horizontal to the vertical force is known as the ‘utilised coefficient of friction’ and slippage will occur if this exceeds the actual coefficient of friction between the foot and the ground. In normal walking, a coefficient of friction of 0.35–0.40 is generally sufficient to prevent slippage; the most hazardous part of the gait cycle for slippage is initial contact. There is a fairly extensive literature on foot-to-ground friction and slippage, e.g. Cham and Redfern (2002) and Burnfield et al. (2005).
The cadence is the number of steps taken in a given time, the usual unit being steps per minute. In most other types of scientific measurement, complete cycles are counted, but as there are two steps in a single gait cycle, the cadence is a measure of half-cycles. The normal ranges for both cadence and cycle time in both sexes at different ages are showed in Table 1 at the end of this chapter, where we consider the effect of age in more detail.
The speed of walking is the distance covered by the whole body in a given time. It should be measured in metres per second. Many authors use the term ‘velocity’ in place of ‘speed’ but this is an incorrect usage of the term, unless the direction of walking is also stated, since velocity is a vector. The instantaneous speed varies from one instant to another during the walking cycle, but the average speed is the product of the cadence and the stride length, providing appropriate units are used. The cadence, in steps per minute, corresponds to half-strides per 60 seconds or full strides per 120 seconds. The speed can thus be calculated from cadence and stride length using the formula:
The walking speed thus depends on the two step lengths, which in turn depend to a large extent on the duration of the swing phase on each side. The step length is the amount by which the foot can be moved forwards during the swing phase, so that a short swing phase on one side will generally reduce the step length on that side. If the foot catches on the ground, this may terminate the swing phase and thereby further reduce both step length and walking speed. In pathological gait, the step length is often shortened, but it behaves in a way which is counterintuitive. When pathology affects one foot more than the other, an individual will usually try to spend a shorter time on the ‘bad’ foot and correspondingly longer on the ‘good’ one. Shortening the stance phase on the ‘bad’ foot means bringing the ‘good’ foot to the ground sooner, thereby shortening both the duration of the swing phase and the step length on that side. Thus, a short step length on one side generally means problems with single support on the other side.
When making comparisons between individuals, particularly with children, it is useful to allow for differences in size. This is done by dividing a measurement by some aspect of body size, such as height (stature) or leg length, a procedure generally known as ‘normalisation’. It is thus fairly common to see walking speed expressed in ‘statures per second’ or to see measures such as ‘step factor’, which is step length divided by leg length (Sutherland, 1997).
Since walking speed depends on both cadence and stride length, it follows that speed may be changed by altering only one of these variables, for instance by increasing the cadence while keeping the stride length constant. In practice, however, people normally change their walking speed by adjusting both cadence and stride length. Sekiya and Nagasaki (1998) defined the ‘walk ratio’ as step length (m) divided by step rate (steps/min) and found that it was fairly constant in both males and females over a range of walking speeds from very slow to very fast. Macellari et al. (1999) made a detailed study of the relationships between gender, body size, walking speed, gait timing and foot placement.
The purpose of this section is to provide the reader with an overview of the gait cycle, to make the detailed description which follows a little easier to follow. The cycle is illustrated by Figures 2.4 and 2.10–2.18, all of which are taken from a single walk by a 22-year-old normal female, weight 540 N (55 kg, 121 lb), walking barefoot with a cycle time of 0.88 s (cadence 136 steps/min), a stride length of 1.50 m and a speed of 1.70 m/s. The individual measurements from this subject do not always correspond to ‘average’ values, because of the normal variability between individuals, although they are all close to the normal range. The measurements were all made in the plane of progression, which is a vertical plane aligned to the direction of the walk; in normal walking it closely corresponds to the sagittal plane of the body. The data were obtained using a Vicon motion system and a Bertec force platform. It should be noted that different laboratories use different methods of measurement, so that other publications may quote different values for some of the measured variables. The reader should thus concentrate on the changes in the variables during the gait cycle, rather than on their absolute values.
Fig. 2.10 • Typical activity of major muscle groups during the gait cycle. Abbreviations as in Figure 2.5. The timings of the events of the gait cycle are typical and not derived from a single subject.
Fig. 2.14 • (A) Mid-stance: position of right leg (green), left leg (grey) and ground reaction force vector 100 ms after opposite toe off. (B) Mid-stance event when the anterior posterior component of the ground reaction force is zero.
When examining diagrams of the joint angles through the gait cycle, it is essential to understand how the angles are defined. Generally speaking, the knee angle is defined as the angle between the femur and the tibia and there is usually no ambiguity. The ankle angle is usually defined as the angle between the tibia and an arbitrary line in the foot. Although this angle is normally around 90°, it is conventional to define it as 0°, dorsiflexion and plantarflexion being movements in the positive and negative directions. In this book, dorsiflexion is a positive angle, but in some other publications it is negative. The ‘hip’ angle may be measured in two different ways: the angle between the vertical and the femur, and the angle between the pelvis and the femur. The latter is the ‘true’ hip angle and is usually defined so that 0° is close to the hip angle in the standing position. Forward flexion of the trunk appears as hip flexion when the hip angle is defined with reference to the pelvis, but not when it is defined with reference to the vertical.
The descriptions which follow assume that symmetry is present between the two sides of the body. This is approximately true for normal individuals, although detailed examination shows that everyone has some degree of asymmetry (Sadeghi, 2003). Such subtle asymmetries are negligible, however, when contrasted with the majority of pathological gaits.
Some gait studies are performed barefoot and some with the subject wearing shoes. Oeffinger et al. (1999) found small differences in some of the gait parameters between these two conditions in children, but did not consider them to be clinically significant. It is usually at the discretion of the investigator whether or not shoes are worn, although in some cases (e.g. when an ankle–foot orthosis or an orthotic insole is used) this may be dictated by the subject’s condition.
During gait, important movements occur in all three planes – sagittal, frontal and transverse. However, this introductory text will concentrate on the sagittal plane, in which the largest movements occur. For information on the motion in other planes, the reader is referred to more detailed texts, such as Perry (1992), Inman et al. (1981) or Rose and Gamble (1994). Figure 2.4 shows the successive positions of the right leg at 40 ms intervals, measured over a single gait cycle. Figure 2.5 shows the corresponding sagittal plane angles at the hip, knee and ankle joints and Figure 2.6 shows the sagittal plane angular velocity of the hip, knee and ankle joints.
Fig. 2.5 • Sagittal plane joint angles (degrees) during a single gait cycle of right hip (flexion positive), knee (flexion positive) and ankle (dorsiflexion positive). IC = initial contact; OT = opposite toe off; HR = heel rise; OI = opposite initial contact; TO = toe off; FA = feet adjacent; TV = tibia vertical.
Figure 2.7 shows the internal joint moments (in newton-metres per kilogram body mass) and Figure 2.8 the joint powers (in watts per kilogram body mass). Different authors have used different units for the measurement of moments and powers; those used here are scaled for body mass, but not for the length of the limb segments. In Figure 2.8, the annotations H1–H3, K1–K4 and A1–A2 refer to the peaks of power absorption and generation described by Winter (1991).
Fig. 2.7 • Sagittal plane internal joint moments (newton-metres per kilogram body mass) during a single gait cycle of right hip (extensor moment positive), knee (extensor moment positive) and ankle (plantarflexor moment positive). Abbreviations as in Figure 2.5.
Fig. 2.8 • Sagittal plane joint powers (watts per kilogram body mass) during a single gait cycle of right hip, knee and ankle. Power generation is positive, absorption is negative. See text for meaning of H1, H2, etc. Other abbreviations as in Figure 2.5.
Figure 2.9 shows a ‘butterfly diagram’, described by Pedotti (1977). This is a plot of the ground reaction vectors and is made up of successive representations, at 10 ms intervals, of the magnitude, direction and point of application of the ground reaction force vector. The vectors move across the diagram from left to right and create a shape that resembles the wings of a butterfly.
Figure 2.10 gives the typical activity of a number of key muscles or muscle groups during the gait cycle. It is based largely on data from Perry (1992), Inman et al. (1981) and Rose and Gamble (1994). Similar, though not identical, data for these and other muscles were given by Sutherland (1984) and Winter (1991). Although Figure 2.10 shows a typical pattern, it is not the only possible one. One of the interesting things about gait is the way in which the same movement may be achieved in a number of different ways and this particularly applies to the use of muscles, so that two people may walk with the same ‘normal’ gait pattern but using different combinations of muscles. The pattern of muscle usage not only varies from one subject to another but it is also affected by fatigue and varies with walking speed, in a single person. The muscular system is said to possess ‘redundancy’, which means that if a particular muscle cannot be used, its functions may be taken over another muscle or group of muscles. A good review of muscle activity in gait was provided by Shiavi (1985).
Figures 2.11–2.19 show the positions of the two legs and the ground reaction force vector beneath the right foot (where present), at the seven major events of the gait cycle and at two additional points – near the beginning of the loading response (Fig. 2.12) and halfway through mid-stance (Fig. 2.14). The description is based on a gait cycle from right initial contact to the next right initial contact. However, the gait cycle could just as easily have been defined using the left leg.
Throughout the text, references will be made to the position of the ground reaction force vector relative to the axis of a joint and to the resulting joint moments. This approach, known as ‘vector projection’, is an approximation at best, since it neglects the mass of the leg below the joint in question (especially important at the hip) and also ignores the acceleration and deceleration of the limb segments (which primarily lead to errors in the swing phase). However, the author has used this approach since it makes it much easier to understand joint moments. The graphs for joint moments (Fig. 2.7) and joint powers (Fig. 2.8) were calculated ‘correctly’, using a method known as ‘inverse dynamics’, which is based on the kinematics, the ground reaction force and the subject’s anthropometry. Wells (1981) discussed the relative merits of these two methods for estimating joint moments.
The upper body moves forwards throughout the gait cycle. Its speed varies a little, being fastest during the double support phases and slowest in the middle of the stance and swing phases. The trunk twists about a vertical axis, the shoulder girdle rotating in the opposite direction to the pelvis. The arms swing out of phase with the legs, so that the left leg and the left side of the pelvis move forwards at the same time as the right arm and the right side of the shoulder girdle. Lamoth et al. (2002) made a detailed study of the relative motion between the pelvis and the trunk at different walking speeds. Murray (1967) found average total excursions of 7° for the shoulder girdle and 12° for the pelvis, in adult males walking at free speed. The fluidity and efficiency of walking depend to some extent on the motions of the trunk and arms, but these movements are commonly ignored in clinical gait analysis and have been relatively neglected in gait research. The whole trunk rises and falls twice during the cycle, through a total range of about 46 mm (Perry, 1992), being lowest during double support and highest in the middle of the stance and swing phases. An approximation to this vertical motion can be seen in the position of the hip joint in Figure 2.4. The trunk also moves from side to side, once in each cycle, the trunk being over each leg during its stance phase, as might be expected from the need for support. The total range of side-to-side movement is also about 46 mm (Perry, 1992). The pelvis, as well as twisting about a vertical axis, also tips slightly, both backwards and forwards (with an associated change in lumbar lordosis) and from side to side. The spinal muscles are selectively activated so that the head moves less than the pelvis, which is important for providing a stable platform for vision (Prince et al., 1994).
The hip flexes and extends once during the cycle (Fig. 2.5). The limit of flexion is reached around the middle of the swing phase and the hip is then kept flexed until initial contact. The peak extension is reached before the end of the stance phase, after which the hip begins to flex again.
The knee shows two flexion and two extension peaks during each gait cycle. It is more or less fully extended before initial contact, flexes during the loading response and the early part of mid-stance (‘stance phase knee flexion’), extends again during the later part of mid-stance, then starts flexing again, reaching a peak during initial swing (‘swing phase knee flexion’). It extends again prior to the next initial contact.
The ankle is usually within a few degrees of the neutral position for dorsiflexion/plantarflexion at the time of initial contact. After initial contact, the ankle plantarflexes, bringing the forefoot down onto the ground. During mid-stance, the tibia moves forward over the foot, and the ankle joint becomes dorsiflexed. Before opposite initial contact, the ankle angle again changes, a major plantarflexion taking place until just after toe off. During the swing phase, the ankle moves back into dorsiflexion until the forefoot has cleared the ground (around feet adjacent), after which something close to the neutral position is maintained until the next initial contact. In the frontal plane, the foot is slightly inverted (supinated, adducted or varus) at initial contact. The foot pronates as it contacts the ground, then moves back into supination as the ankle angle changes from plantarflexion to dorsiflexion, this supinated attitude being maintained as the heel rises and the ankle plantarflexes prior to toe off. Some degree of supination is retained throughout the swing phase.
Each of the following sections begins with some general remarks about the events surrounding a particular event in the gait cycle and then describes what is happening in the upper body, hips, knees, ankles and feet, with particular reference to the activity of the muscles. These sections are very detailed and may be too much to comprehend in one ‘pass’. It is suggested that the reader should skip the moments and powers on the first reading, but should go back to them later, to gain a deeper understanding of the mechanical processes underlying the gait cycle. The figures shown in this section represent the normal positions of the lower limbs and pelvis at different events during gait and the ground reaction force vector expected.
Initial contact is the beginning of the loading response, which is the first period of the stance phase. Initial contact is frequently called ‘heelstrike’, since in normal individuals there is often a distinct impact between the heel and the ground, known as the ‘heelstrike transient’. Other names for this event are ‘heel contact’, ‘footstrike’ or ‘foot contact’. The direction of the ground reaction force changes from generally upwards during the heelstrike transient (Fig. 2.11) to upwards and backwards in the loading response, immediately afterwards (Fig. 2.12). This change in direction can also be seen in the butterfly diagram (Fig. 2.9), where the force vector changes direction immediately after initial contact.
The trunk is about half a stride length behind the leading (right) foot at the time of initial contact. In the side-to-side direction, the trunk is crossing the midline in its range of travel, moving towards the right, as the foot on that side makes contact. The trunk is twisted, the left shoulder and the right side of the pelvis each being at their furthest forwards and the left arm at its most advanced. The amount of arm swinging varies greatly from one person to another and it also increases with the speed of walking. At the time of initial contact, Murray (1967) found the mean elbow flexion was 8° and the shoulder flexion 45°.
The attitude of the legs at the time of initial contact is shown in Fig. 2.11. The maximum flexion of the hip (generally around 30°) is reached around the middle of the swing phase, after which it changes little until initial contact. The hamstrings are active during the latter part of the swing phase (since they act to prevent knee hyperextension); gluteus maximus begins to contract around the time of initial contact and together these muscles start the extension of the hip, which will be complete around the time of opposite initial contact (Fig. 2.5).
The knee extends rapidly at the end of the swing phase, becoming more or less straight just before initial contact and then starting to flex again (Figs 2.5 and 2.11). This extension is generally thought to be passive, although Perry (1992) states that it involves quadriceps contraction. Except in very slow walking, the hamstrings contract eccentrically at the end of the swing phase, to act as a braking mechanism to prevent knee hyperextension. This contraction continues into the beginning of the stance phase.