Free joint mobility and appropriate muscle force increase walking efficiency. As the body moves forward, one limb typically provides support while the other limb is advanced in preparation for its role as the support limb. The gait cycle (GC), in its simplest form, consists of stance and swing phases. The stance phase is subdivided into three segments, including (1) initial double stance, (2) single limb stance, and (3) terminal double limb stance.
Each double stance period accounts for 10% of the GC, while single stance typically represents 40% (60% total). The 2 limbs typically do not share the load equally during double stance periods. The swing phase for this same limb is the remaining 40% of the GC. Ipsilateral swing temporally corresponds to single stance by the contralateral limb. Slight variations occur in the percentage of stance and swing related to gait velocity. Duration of each aspect of stance decreases as walking velocity increases. The transition from walking to running is marked by elimination of double support period(s).
A stride is the equivalent of a GC. The duration of a stride is the interval between sequential initial floor contacts by the same limb. A step is recognized as the interval between sequential floor contacts by ipsilateral and contralateral limbs. Two steps make up each GC, which is roughly symmetric in normal individuals.
A consistent sequence of motions is performed at each of the lower extremity joints during locomotion. Each stride contains 8 relevant phases. Stance consists of 5 gait phases (ie, initial contact, loading response, midstance, terminal stance, preswing), with the remaining 3 phases occurring during swing.
The first 2 gait phases (0-10% GC) occur during initial double support. These phases include initial contact and the loading response. Initial contact often is referred to as heel strike. While this term is appropriate in normal gait, many patients achieve heel contact later in the GC, if at all. The joint motion during this phase allows the transfer of weight onto the new stance phase leg while attenuating shock, preserving gait velocity, and maintaining stability.
Swing phase by the contralateral limb corresponds with single support by the ipsilateral limb to support body weight in the sagittal and coronal planes. The first half of single support is termed midstance (10-30% GC) and is involved with progression of the body center of mass over the support foot. This trend continues through terminal stance (30-50% GC). This phase includes heel rise of the support foot and terminates with contralateral foot contact. [1]
The final stance element, preswing (50-60% GC), is related functionally more to the swing phase that follows than to the preceding stance phase events. Preswing begins with terminal double support and ends with toe-off of the ipsilateral limb.
Three unique phases characterize swing, including initial swing (60-73% GC), mid swing (73-87% GC), and terminal swing (87-100% GC). The swing phase achieves foot clearance and advancing of the trailing limb.
Shock absorption and energy conservation are important aspects of efficient gait. Altered joint motion or absent muscle forces may increase joint reaction (contact) forces and lead subsequently to additional pathology. In early stance, nearly 60% of one's body weight is loaded abruptly (less than 20 milliseconds) onto the ipsilateral limb. This abrupt impact is attenuated at each of the lower extremity joints. Loading response plantar flexion is passive, substantially restrained by eccentric work of pretibial muscles. The absorptive work by pretibial muscles delays forefoot contact until late in the initial double support period (7-8% GC).
At initial contact, external (ground reaction) forces applied to the contact foot produce a tendency toward knee flexion. Repositioning the knee (recurvatum) increases knee mechanical stability, but at the cost of increased contact forces and shock generation. A balance between knee stability and shock absorption is achieved by eccentric quadriceps contractions during loading response. The impact of loading is minimized at the hip during single support through hip abductor muscle contraction. [1]
A study by Ricci et al demonstrated the importance of the pubic rami in maintaining pelvic integrity during loading in the gait cycle. Evaluating pubic ramus fractures, the investigators found that if the anterior pelvic ring rami are completely disrupted, load redirection causes significantly greater posterior pelvic stress. [2]
Ambulation always is associated with metabolic costs. These costs are relatively minor in normal adults performing free speed level walking. The self-selected walking speed in normal adults closely matches the velocity that minimizes metabolic work. This association does not apply with gait pathology. Walking velocity, energy cost per time, and energy cost per distance are considerations when the patient is making choices about walking versus wheelchair mobility. Gait velocity typically decreases with all neuromuscular pathology, and the reduction is related to the magnitude of the pathology. Energy cost per unit of time may not change substantially, even with severe involvement. Energy cost per unit of time is maintained by decreasing walking velocity considerably. Energy cost per unit of time does not change markedly following stroke, as compared to changes associated with aging; however, the energy requirement per distance traveled is more than 3 times normal.
In this same population, wheelchair use cuts energy cost per distance in half and decreases cost per minute slightly, while preserving ambulation velocity. Similar trends are observed when examining various energy cost parameters in individuals with spinal cord injury, myelomeningocele, and increasing levels of amputation. Energy cost to travel a prescribed distance increases (greater than 500% increase in myelomeningocele with bilateral knee-ankle-foot orthoses), while oxygen cost per minute is maintained by decreasing walking velocity substantially. Often the critical factor in selecting a wheelchair for mobility is the energy requirement to traverse a given distance. Most individuals self-select wheelchair mobility when cost per distance exceeds 300% of normal values.
Most commonly, observational gait analysis is appropriate to characterize most gait pathologies. This approach is sufficient to note gross abnormalities in walking; however, as walking complexity increases with organic pathology, objective analysis becomes necessary. The measurement systems that follow may be used individually or collectively. The choice of method is based on clinical need, financial considerations, and staffing at the specific laboratory. Typically, motion and force data are compiled simultaneously. [3]
Instrumented motion analysis is a logical extension of observational gait analysis. The level of complexity involved in examining 3 joints each in 2 limbs is such that few individuals can perform this task consistently. Measurement of human motion is complex. Although most gait motion occurs in the sagittal plane, subtle rotations in the other planes are clinically important. Analysis of motion in the sagittal plane is confounded if that motion does not lie within the plane recorded by a single camera. Rotations of the limb are attenuated when the limb is internally or externally rotated with respect to the camera's plane. Three-dimensional motion analysis overcomes this drawback.
Position of joint centers of rotation is estimated by markers placed on the skin surface. Markers may be enhanced with passive markers or active light-emitting diodes. Motion of the limb segments about the joint centers is recorded. At least 3 markers are required per limb segment, and the position of each marker is recorded using multiple cameras. A calibration system is used to translate film to real-life dimensions. Angular position of each segment may be determined for each percent GC interval. Angular velocity and acceleration of the limb segments are obtained by mathematical differentiation and smoothing of limb position data. Motion data may be combined with analysis of external forces that act on the body.
An alternative technique to quantify motion is through use of electrogoniometers. A triaxial electrogoniometer has parallelograms and potentiometers for the sagittal, coronal, and transverse planes. Electrogoniometers have several limitations, but the advantages of this method include convenience, ease of use, immediate availability of data, and decreased expense when compared to video acquisition techniques.
Calculation of joint moments (torque) and reaction forces between segments is dependent on knowledge of the inertial components of the respective segments (kinematics), [1] body segment parameters, and external forces (kinetics) that affect the body.
The magnitude and distribution of segment masses with respect to joint axes are obtained from cadaveric data or mathematical modeling of limb segments. The contact (ground reaction) force applied at the distal segment is measured with a force platform. This thin plate typically measures forces and moments in 3 dimensions about the foot center of pressure. Oscillation of the center of pressure may be obtained in the same manner.
Integration of external force, center of pressure, unique body segment parameters, and motion data yields information on joint moments, joint power, and reaction forces between segments using standard inverse dynamics techniques. The role of muscle groups is inferred from the magnitude and sign of the moments and power at the respective lower extremity joints.
Muscle action cannot be measured directly. The electromyogram (EMG) allows indirect measurement of muscle activity. While single motor units are analyzed routinely in clinical electrodiagnostic studies, this level of analysis is not performed as part of a routine gait analysis. A typical kinesiologic (dynamic) EMG represents the activity of multiple motor units. The stochastic nature of the waveform confounds simple analysis. The interference pattern recorded demonstrates both mechanisms used to increase force output and increase in the number and in the firing rate of motor units.
Kinesiologic EMG contains important information about the timing and relative intensity of the signal. Timing of the signal is straightforward. Muscle electrical activity precedes force generation by 40-120 milliseconds. This electromechanical delay is associated with compliance of the tendon at the onset of the EMG signal and continued binding of actin-myosin crossbridges after termination of neural drive related to delays in calcium resequestration.
Muscle force cannot be estimated directly from the relative intensity of the signal. A linear relationship between EMG intensity and force output has been demonstrated only in isometric contractions. Presentation of the quantified EMG may be in absolute values (volts) or as a percentage of some normalized standard. Absolute voltage does not represent clinically significant effort, as the electrode type and placement substantially affect signal magnitude. Normalization of muscle force output allows for comparisons across muscles related to relative intensity. Gait EMG is expressed as a percentage of maximum muscle contraction. Another technique is to express gait EMG output as a percentage of gait maximum. This latter technique often must be used in persons who cannot give a maximum voluntary contraction as a result of decreased neural control. A disadvantage is that peak values of weak and strong muscle forces are defined as 100%.
Mechanical work is the integral of force and velocity over time and the product of joint power and joint angular velocity. Mechanical work includes (1) changes in the mechanical energy of the body links; (2) energy expenditure for mechanical movement as the sum of the positive values, including general center of mass energy changes (external work); and (3) changes in energy in the links in their movement relative to the center of mass (internal work).
Limb movement requires energy for muscle contraction. Mechanical work does not equal metabolic work in most instances. For example, mechanical work in a cyclic motion with constant average speed (without dissipation of energy) is equal to 0 since the mechanical energy of the system has not changed (net 0 paradox). Metabolic energy is required for mechanical movement in this case. Energy metabolism is another means of assessing the cost of locomotion.
Energy expenditure may be determined by indirect calorimetry; however, in most later literature, the O2 gas volume is reported without conversion to calories. The volume of oxygen consumption per body weight enables intersubject comparisons and provides information on overall gait performance. Measurement of oxygen consumption (metabolic energy expenditure, VO2) generally is reported under standard conditions of temperature (0°C), pressure (760 mm Hg), and dry conditions (no water vapor). After 2-3 minutes of exercise at a constant submaximal workload, the rate of oxygen consumption reaches a level sufficient to meet tissue demand for oxygen. Parameters of physical work (cardiac output, heart rate, and respiratory rate) reach a steady state. Energy consumption for the activity at this time is reflected by the rate of oxygen consumption. The O2 rate determines the intensity and duration the exercise can be performed.
The rate of oxygen consumption is lowest at comfortable walking speeds in normal adults (approximately one third of maximal aerobic capacity). Slight increases in energy demands are demonstrated at slower walking speeds; more substantial increases are observed as walking speed increases above the upper limits of normal. Energy expenditure at each walking speed is not related to whether the patient is male or female.
The arcs of motion at the ankle are relatively small; yet, they are essential for shock absorption and progression of the body's center of mass. The ankle plantar flexes throughout loading response. Dorsiflexion begins with single support, as the tibia rotates forward over the fixed foot. Rapid plantar flexion begins at terminal double support, with maximum plantar flexor position of 30° attained at toe-off. This action marks the initiation of swing with dorsiflexion throughout the 3 swing-phase segments.
Motor control at the ankle is understood most easily by commencing with swing phase muscle activity. Ankle dorsiflexors undergo brief eccentric contraction in preswing, followed shortly by concentric contraction at the initiation of swing. This contraction sequence increases mechanical efficiency and assures foot clearance. Continued pretibial muscle isometric contraction through swing maintains this neutral or slightly dorsiflexed posture. Subsequent loading response is notable for eccentric contraction of the pretibial muscles to control plantar flexion. A brief period of co-contraction of ankle plantar flexors and dorsiflexors occurs at the transition from initial stance double support to single support. This co-contraction interval increases limb stability and may smooth the transition from double to single support.
The orientation of the ankle and subtalar axes couples loading response dorsiflexion with eversion. Both are attenuated by eccentric contraction of the ankle inverters (ie, posterior tibialis). The plantar flexors begin force generation in single support, with peak activity late in terminal stance and preswing. The absorptive power demonstrated in single support is related to their role in restraining forward tibial rotation. The timing of the triceps surae and perimalleolar muscles is similar to that of the triceps surae (gastrocnemius plus soleus) that supplies most of the ankle plantar flexor moment. The ankle then is plantar flexed vigorously during preswing.
The role of the calf muscles during preswing is controversial. Some authors describe this gait phase as push-off and suggest that muscle force by the ipsilateral calf actively propels the limb (and body center of mass) forward. An alternative model is proposed by Perry, who states that the ipsilateral limb makes no active contribution to the push-off. A third model is that active plantar flexion occurs at the ankle (positive power) and does not serve to propel the body forward (no increase in walking velocity), but, instead, functions as part of a closed kinetic chain to initiate knee and hip flexion in preparation for swing. This phenomenon may account for the increased knee and hip power requirements in individuals who do not plantar flex actively in preswing (ie, transtibial amputees).
Most knee motion is limited to the sagittal plane. The knee travels from slight knee flexion at initial contact (5°) to nearly 20° of flexion by the end of loading response. The knee then extends (net flexion) through single support, with peak stance phase extension at 40% GC. At the conclusion of terminal stance and preswing, knee flexion is rapid, continuing through initial swing until peak knee flexion (60°) occurs. This trend then is reversed, with knee extension continuing through terminal swing. Peak knee extension occurs slightly before the end of the swing phase, with minor flexion occurring in preparation for the subsequent stance phase.
The role of knee muscles, like that of ankle muscles, is understood most easily if one begins the analysis as the limb is in swing. In early swing phase, knee flexion is passive, resulting from active plantar flexion and hip flexion. During swing, motion at the hip changes from flexion to extension, with the knee passively extended as a result of this change of hip joint direction. All 3 hamstrings (long head of the biceps femoris, semimembranosus, semitendinosus) are active in mid-to-late swing to decelerate the extending knee. The quadriceps is also active in terminal swing. This co-contraction of knee flexors and extensors prepares the limb for weight acceptance that follows shortly. The quadriceps continues to be active in loading support. The eccentric activity of these muscles attenuates the shock of weight acceptance, while preventing excess knee flexion (< 20°). [4]
Both the hamstring and quadriceps muscles are quiescent during midstance while the knee is extended passively. The role of the remaining quadriceps muscle is different. The rectus femoris is active electrically for only a brief period corresponding with preswing and initial swing. Although this muscle crosses and extends the knee, its role in gait is to assist in flexion at the hip as swing commences. [5]
The gluteus maximus timing and relative intensity are similar to those demonstrated by the other hip extensors (eg, hamstrings). During late swing, the gluteus maximus functions to reverse hip flexion to extension. The gluteus maximus, generally the body's single strongest muscle, resists external forces from loading response that would flex the hip excessively (eg, jack-knifing). A similar EMG profile and role are demonstrated by the adductor magnus. Unlike in the knee and ankle, there is a substantial amount of frontal plane motion at the hip. During single support, the mass of the torso tends to rotate the body about the stance limb hip joint axis (eg, contralateral lateral tilt). Typically, this motion is limited by eccentric activity of the gluteus medius, gluteus minimus, and, to a lesser extent, the tensor fascia lata.
Hip flexion marks transition from stance to swing. This motion advances the trailing limb and assists in foot clearance, which is accomplished by concentric contractions of the iliopsoas, rectus femoris, and sartorius muscles. Swing phase activity of the hip adductors brings the feet toward the line of progression, decreasing the energy demands of walking.