Assessing gaits in older adults for the purposes of preventing falls, decreasing the risks of falls, and minimizing injuries from falls is a complex task. This complexity can be ascribed to the involvement of multiple body systems and requires a multipronged approach to understanding the effects of these systems on age-associated gait changes. Comprehend these effects is difficult without first delving into the mechanics of the elderly gait. Here we present a simpler approach to recognizing some of these multifaceted changes for practitioners and researchers alike. With advancing age, there are modifications in spatiotemporal, kinematic, and kinetic gait parameters that could increase the risk of falling. These age-associated changes in gait characteristics can be classified as changes in both propulsion and stability. A decrease in gait velocity as a result of less forceful ankle plantar flexor torque and greater hip contribution not only increases the likelihood of a fall but also alters the fall direction, which in itself increases the chances of a detrimental injury. Understanding the foot to floor relationship is vital because older adults tend to walk with more plantarflexion when compared to young adults. This increases the chance of tripping either by stubbing the toe during the swing phase when the foot’s trajectory velocity is the greatest or catching the bottom of the shoe immediately prior to foot contact. Older adults tend to produce increased lateral movement in relation to their base of support when compared to young adults, thus affecting their stability. These observable characteristics of the elderly gait can be attributed to the underlying structural and physiological changes in the neuromuscular system. Therefore, the contributing factors of proprioception, vision, and vestibular system, as well as aging muscle, brain, and spinal cord, will be addressed to determine their contribution to a fall.
Elderly gait mechanics; elderly fall risk; spatiotemporal gait parameters; kinematic gait parameters; kinetic gait parameters; gait mobility; aging gait stability; physiological aging gait factors
Human locomotion and gait occur as a result of complex interplays of multiple muscle contractions across multiple joints as affected by different environmental conditions (internal and external) in which these contractions occur. For successful locomotion, we must consider not only that a person has traversed a distance from one point to another but also factors such as pain, fatigue, balance, safety, and quality of movement. Age-associated declines in body functions can affect most of these factors in unison or in various combinations. Failure to maintain an effective gait could be a result of any of these factors or their underlying causes, which might disrupt activities of daily living and result in increased fall risk, especially in older adults.
Successfully completing daily living activities and maintaining independence are closely associated with quality of life, and a healthy gait is essential for both. In this regard, gait analysis and interpreting gait adaptations in older adults as a function of health status and falls has been of prime importance for more than 20 years for physical therapists, biomechanists, and physicians. In this chapter, we will focus on age-induced changes in gait on level surfaces along with changes in the neuromuscular system that may result in alterations that predispose older adults to increased risk of falls.
The chapter is divided into two broad sections: (1) mechanics of gait changes and (2) changes in physiological functions. In the first section, changes in gait in older adults will be approached from two primary directions: changes in mobility and changes in stability. Changes in mobility will address biomechanical parameters of forward progression. Changes in stability will explain factors that affect older adults’ gait stability using conventional biomechanical variables of stability during gait. The second section will incorporate changes in balance and how various systems integral to maintaining balance change with age. This will be followed by changes in functioning of the core gait apparatus: the human neuromuscular system.
Gait is a function of the entire body and is directly and indirectly affected by almost everybody system. Here we will be talking strictly about aging-induced changes in the neuromuscular system and their effects on gait. These neuromuscular processes involve changes at muscle level, motor neurons, and the central nervous system (CNS). Changes in gait with aging can be attributed to pathology or healthy aging or both, so it is important while reviewing literature that demographic characteristics of population tested are well described and understood. One primary reason for such elucidation is that healthy older adults living an active lifestyle maintain gait changes that are similar to young adults well into old age (Bloem et al., 1992; Verghese et al., 2006). Some of the changes in gait such as slowness, increased step width, and stooped posture can be attributed to subclinical pathologies that may affect overall quality of life and independence in older adults. Regardless, certain changes in gait do occur with increasing old age. Some of these changes are irreversible but others can be arrested and their damages minimized through exercises and physical therapy.
As people age, the speed they prefer to walk without restriction significantly declines. Reasons for the decline in this preferred gait velocity include reduction in lower-body muscle strength (Bassey et al., 1992), aerobic capacity (Mian et al., 2006), cognitive function status (Buchner et al., 1996), and physical health (Cesari et al., 2006). Mechanically, reduced gait speed is associated with reduced step length, an increase in the double-stance phase, and gait stability ratio.
Murray and coauthors studied gait parameters for 64 men ranging in age from 20 to 87 years of age and reported that the oldest three age groups—67–73, 74–80, and 81–87 years of age—produced the slowest preferred gait speed of 1.18, 1.23, and 1.18 m/s and fast walking speeds of 1.63, 1.67, and 1.60 m/s, respectively (Murray et al., 1969), when compared to the other age groups. The younger groups averaged 1.52 and 2.15 m/s for preferred and fast walking speeds (Murray et al., 1969). Stride length was significantly reduced for the oldest three walking groups during both preferred and fast walking speeds at 1.36, 1.41, and 1.26 m and 1.60, 1.59, and 1.40 m, respectively. Therefore, males younger than 65 years of age averaged stride lengths that were approximately 89% of their height; for those 67 and older, the percentage drops to 79% of height for preferred walking speed. At fast gait speeds, the percentages climbed to 107% and 90% for the younger groups in comparison to the three older groups (Murray et al., 1969).
Elble and colleagues compared gait parameters for 20 young (30.0 ±6.1 years) and 20 older adults (74.7 ±6.6 years) who walked a 10-m walkway four times at their preferred and fast walking speeds (Elble et al., 1991). The preferred and fast gait speeds were 1.18 and 1.67 m/s and 0.94 and 1.39 m/s, respectively for young and older adults, which equated to a 20% and 17% decline for the older adults in comparison to the young adults. The older adults had 1.08 and 1.26-m stride lengths for preferred and fast walking speed in comparison to young adults at 1.32 and 1.58 m. However, these changes were not due to cadence because cadence was reported to be similar at 107 and 104 steps/min and 126 and 128 steps/min for young and older adults at preferred and fast walking speeds (Elble et al., 1991). These results indicate that healthy older adults take shorter steps at a similar frequency resulting in a decrease in gait velocity for both preferred and fast gait speeds. In addition, the shorter steps taken by the older adults increase the amount of time spent in stance and the double-support phase (Elble et al., 1991).
Nigg and Skleryk (1988) determined that older adults reduced their preferred walking speed either as a result of increased joint stiffness or because of a need to increase safety and balance (Nigg and Skleryk, 1988). This study alludes to an important aspect of age-associated adaptation: whether the changes in gait are a result of a pathology or a need to increase safety. In either situation, the physiological or pathological changes in the human body need to be understood before any judgment can be passed on gait in older adults. This idea has been well explained in an excellent review by McGibbon (2003). As correctly pointed out in that paper, the clarity between adaptations to a primary pathology and changes related to aging itself needs to be explained further. The article argues that most of the changes in gait due to age are a manifestation of a primary pathology.
One prominent change has been a shift to a hip-dominant strategy from an ankle-dominant strategy for forward propulsion (DeVita and Hortobagyi, 2000; Franz and Kram, 2014; Kerrigan et al., 1998). Once again, the shift in strategy is related to the gait speed, and one of the key differences while reviewing various studies arises from whether walking speed was self-selected or not. Judge and colleagues (1996) tested young and older adults at self-selected walking speeds and found a significant reduction in ankle plantar flexor power that was replaced by increased hip flexor power during late stance in older adults. They reported that the primary age-limiting factor in reduction of gait velocity comes from reduction of the ankle plantar flexor push-off. This has been contradicted by other studies that have found that older adults can increase their hip contribution to match young adults at a faster gait velocity but the limitation is in the hip extension range of motion (Kerrigan et al., 1998; Riley et al., 2001; Riley et al., 2001). Riley et al. (2001) analyzed age-associated gait differences and the contribution of lower-extremity joint moment to gait velocity and forward progression. When the elderly were asked to walk at higher velocity, they were able to increase the contribution of ankle plantar flexor power at higher velocity but could not increase their hip contribution, which suggests a limitation in hip extension moments during gait.
DeVita and Hortobagyi (2000) stated that overall support moment is similar between older and young adults, although the moments are redistributed between lower-extremity joints with the hip increasing its role and the ankle decreasing its contribution (DeVita and Hortobagyi, 2000). Increased hip concentric power during early stance and reduced ankle push-off during late stance in old adults suggests this altered gait strategy. Similar results have been reported by other authors comparing various combination of speed (preferred, slow, and fast). The verdict is clear that old adults reduce their ankle plantar flexor push-off and increase the hip concentric and eccentric load in order to continue forward progression. Also, that limitation in an old adult’s ability to generate power through hip joint affects gait velocity more than reduction in ankle moments. However, this last assertion has to be questioned. First, redistribution of support moments occurs as a result of ankles producing lower push-off, therefore changes in hip moments are an adaptation themselves. Second, ankle push-off reduces first, which increases the demand on hip muscles. Third, plantar flexors can still increase their effort if need arises but hip muscles cannot. What needs to be understood is whether the changes in gait strategy is a function of greater or faster rate of strength loss in ankle plantar flexors or is an alteration in joint proprioception and balance. Since the muscle strength was not measured in these studies, it cannot be conclusively said whether these changes are adaptations for weak ankle plantar flexors or are a part of normal aging. Nevertheless, plenty of evidence for reduction in plantar flexor strength in older adults along with other concomitant changes in neuromotor processes is available. Also, changes in hip extensors with aging is also well documented (Dean et al., 2004). Therefore, the loss of strength in hip and ankle muscles cannot by itself explain the change in gait strategy.
On the other hand, the knee joint moments are essential components of support moment required to maintain a dynamic vertical posture during gait, therefore the role of the quadriceps’ muscle strength should be considered a potential mechanism of falling. Quadriceps strength reduces with age similar to other muscles of the lower extremities, but the weakness is associated with failure to maintain upright stance, difficulty in moving from sitting to standing, and instability while going up and down stairs. An indirect link between reduced quadriceps strength and falls can be inferred from the results of a study that found that participants who were obese and had limited knee extensor strength showed a fast decline in their gait velocity, thus predisposing them to increased risk of fall and mortality (White et al., 2013). Quadriceps strength has been considered a significant predictor of falls over a 3-year period in community-dwelling healthy older women (Scott et al., 2014). Quadriceps strength was also a better predictor of first incidence of falls compared to gait velocity. Since ankle and hip muscle strength were not measured, the results cannot be used to compare the relative importance of strength loss in different muscle groups across lower-extremity joints. However, if all the evidence presented here is considered, we need to ask the following questions. Ankle plantar flexors (during high gait speed) and hip extensors can compensate for one weakness in another, but is there a mechanism for compensating for quadriceps strength loss? Mechanically increased activity of hip extensors and ankle plantar flexors can passively extend the knee and provide a rigid lever in the lower extremity. Does that mean the functional loss of quadriceps strength increases the burden on other muscles and is the primary mechanism that predisposes healthy older adults to a fall risk? Research thus far does not provide a clear answer to that question.
Stability and balance are related terms in motor control and biomechanical literature that carry multiple definitions. Clinically, falls occur as a result of loss of balance, whereas stability can be loosely defined as the ability to resist perturbations or forces that result in loss of balance. An analysis of stability during gait can be undertaken using multiple methods, and each method defines stability accordingly. We are going to limit the discussion on stability by utilizing a conventional linear variable called gait stability ratio; kinematic variables such as toe clearance, foot inclination, lateral trunk sway, and heel velocity; and a kinetic variable such as angular momentum. The list is in not all-encompassing because the analysis strategies have differed over the decades and left us an accumulation of variables that can be used to show differences in gait stability among age groups.
Gait stability ratio was first introduced to account for a decrease in preferred walking speed as the ratio of cadence (steps/s) to gait velocity (m/s) and is reported as steps per meter (Cromwell et al., 2001). An increase in the number of steps per unit of distance coupled with a decrease in preferred walking speed is an indication of walking stability, which equates to an increase in the amount of time spent in the double-stance phase (Cromwell and Newton, 2004). An increase in the amount of time spent in the double-stance phase reduces the dynamic component of walking while increasing gait stability (Maki, 1997).
The gait stability ratio for older adults significantly increased when compared to young adults, 1.48 ±0.19 to 1.36 ±0.17 steps/m, respectively (Cromwell and Newton, 2004). The increase in gait stability ratio indicates that older adults take more steps per unit of distance, thereby increasing stability while walking at their preferred speed. An increase in stability while walking allows them to compensate for reductions in balance thus maximizing walking stability, creating resistance to perturbations, and reducing fall potential.
The older adults reduce gait velocity, decrease step length, and increase gait stability ratio, which increases time spent in the double-support phase. Age-related adaptations during gait are considered to be mechanisms that increase stability at the cost of reducing mobility (Cromwell and Newton, 2004); however, this results in individuals becoming less efficient by reducing the forward progression within each gait cycle. The reduced gait velocity brings about a mechanical reduction in the amount of forward progression in a given amount of time and lowering the momentum of the body. This loss of momentum is sometimes considered critical in preventing falls. Conversely, gait patterns among young adults are distinguished by phases of instability, which produces a much more efficient forward progression and therefore a preferred walking speed that is quicker.
The primary focus of fall risk kinematics in the older adults has been tripping, slipping, and lateral movement of the trunk while walking. Tripping and loss of balance remain significant reasons why older adults fall while walking. Tripping accounts for approximately 53% of falls among older adults (Blake et al., 1988) and is separated into two classifications: stubbing the toe and catching the bottom of the foot during the swing phase. A decrease in ankle dorsiflexion during the swing phase, as suggested by Oberg et al. (1994), increases the chances of stubbing the toe along with a greater probability of catching the bottom of the foot. The initial mechanism can be understood by analyzing toe clearance, while the latter can be investigated using the foot’s inclination angle.
Tripping is associated with stubbing the toes on a walking surface or on an obstacle at approximately the midpoint during the swing phase of the gait cycle when the toe velocity is greatest and toes are at their minimal height above the ground (Schulz, 2011; Winter, 1992). The tripping mechanism occurs without an apparent obstacle causing the stumble due to a foot catch at the lowest point of the swing phase or immediately prior to heel strike (Chen et al., 1994). Older adults demonstrate less toe clearance at a point during the swing phase when the velocity of the distal end of the foot is the greatest (Barrett et al., 2010), which increases the chances of catching the toe on an obstacle or the floor. Maximal foot clearance is the minimum vertical clearance between the lowest point of the foot of the swing leg and the walking surface (Barrett et al., 2010), and it is associated with reduced foot inclination during the swing phase of the gait cycle (Chiba et al., 2005).
An increase in lateral sway among older adults while walking is indicated to be a fall variable (Chiba et al., 2005), especially as preferred gait velocity decreases. Lateral sway ratio is the relationship between the center of mass (COM) of the body and the base of support created by the foot placement in the frontal plane (van den Kroonenberg et al., 1996). The displacement of the COM close to or outside the base of support will increase the instability and can potentially result in a fall. Lateral gait unsteadiness has been considered a determining factor in lateral fall risk assessment (Helbostad and Moe-Nilssen, 2003) because the COM displacement in the frontal plane is a source of greater concern because it might result in people falling on their sides. This can potentially cause more debilitating injuries to the pelvis and hip complex and affect long-term outcomes in older adults, including the quality of life.
Chiba and associates compared the differences between fallers and nonfallers among community-dwelling individuals using minimum toe clearance, maximum foot inclination, and the trunk’s lateral sway ratio. Fifty-six older adults were separated into two groups: 25 fallers (76.0 ±6.6 years) and 31 age- and gender-matched nonfallers (74.9 ±7.2 years) who walked a 6-m walkway at their preferred walking speed (Chiba et al., 2005). When compared with the nonfallers, the fallers produced less toe clearance (12.0 ±0.7 mm and 15.2 ±1.0 mm, respectively), lower maximal sole inclination angle (7.4 ±0.8 degrees and 14.3 ±0.9 degrees, respectively), and larger lateral trunk sway (0.23 ±0.01 and 0.18 ±0.01, respectively). This indicates altered ankle–foot dynamics in fallers, which increases the risk of trips while walking on a level smooth terrain. The increase in lateral sway ratio for fallers in comparison to nonfallers indicated a problem with the whole body during the gait, which may be a result of increased lower-extremity variability, reduced muscle strength, and altered reflexes (Chiba et al., 2005). However, none of this was tested and therefore can only be treated as speculation. Variability in trunk movement patterns in older adults demonstrate a different motor control strategy in the mediolateral direction compared to other anteroposterior and vertical directions (Moe-Buksseb and Helbostad, 2005). Therefore, it is essential to test gait stability in both the anteroposterior and mediolateral directions.
Slips are another important cause of falls (Lloyd and Stevenson, 1992). Just like trips, slips can be analyzed using a multitude of variables. The one we are explaining in detail is heel velocity and how it indicates increased risk of falls in older adults, especially as the foot approaches heel contact. The heel velocity immediately prior to heel strike is an indication of foot trajectory during the swing phase and the end-point control (Winter, 1992), so higher heel velocity indicates an increased risk for slipping (Winter et al., 1990). When compared to young adults, the older adults produced significantly higher heel velocity (1.15 and 0.87 m/s, respectively) immediately prior to heel strike (Winter et al., 1990; Winter, 1992), thus indicating greater risk for slipping in older adults even though their preferred walking speed was significantly less (1.29 ±m/s compared to 1.43 ±m/s) than young adults (Winter et al., 1990).
Since gait is a combination of angular movements occurring at the lower extremity and the upper-extremity joint, their combined movement generates angular momentum. Failure to control the angular momentum of the body can result in instability and therefore predispose an individual to an increased risk of falls. Whole-body angular momentum is considered to be a highly controlled variable (Bennett et al., 2010; Neptune and McGowan, 2011) during different human tasks. It may be important in maintaining balance and stability from sitting to standing (Riley et al., 1997) and in trip events (Pijnappels et al., 2004). It has been suggested that angular momentum while walking is controlled by the CNS and the control synergies emanating from it. Thus, the altered regulation of the angular momentum may be a sign of an increased risk of fall (Popovic et al., 2004). Altered patterns of angular momentum regulation has been reported between older men and women while negotiating stairs (Singhal et al., 2015), which highlights the increased risk of falls in older women. However, more studies need to be undertaken to understand age-associated differences in the regulation of angular momentum.
Analysis of balance in itself is a complicated task because of the multiple body systems involved: visual, vestibular, and proprioceptive. Obvious changes in vision due to glaucoma, macular degeneration, diabetic retinopathy, and bifocal lenses can lead to a loss of balance and are relatively easier to comprehend. However, subtle changes in depth perception, ground overlay, three-dimensional forms, and slant characteristics are much more difficult to discern but have been found to be reduced in older adults. All these changes would affect older adults’ ability to navigate natural environment and would either increase their risk of falls or cause adaptations in gait that would allow them to be more unstable while walking. Vestibular causes of imbalance can range from specific pathologies such as benign paroxysmal positional vertigo and Méniére’s disease to idiopathic changes in vestibular apparatus. An age-associated decrease in functional vestibular connectivity and an increase in its variability have been reported in the vestibular cortical network (Cyran et al., 2016). The authors have suggested reduced cortico–cortical inhibition as a potential mechanism for this impairment, which may increase the risk of falls.
Proprioception is a neural correlate of balance and comprises inputs through mechanoreceptors located in the joints and muscles to the CNS. The two components of proprioception—the joint position sense and the sense of limb movement—are both essential for coordinated movement patterns, motor control during posture, and gait and motor learning (Ghez et al., 1995; Ghez and Sainburg, 1995; Hiemstra et al., 2001; Pickard et al., 2003; Tsang and Hui-Chan, 2003). People of all age groups are more dependent on proprioception to maintain balance than vision and vestibular sensation (Colledge et al., 1994). The occurrence of proprioceptive decline with age and its effects on balance and motor control have long been identified (Barrack et al., 1984; Bullock-Saxton et al., 2001; Horak et al., 1989; Kaplan et al., 1985; Lord and Ward, 1994; Manchester et al., 1989; Pai et al., 1997; Petrella et al., 1997; Skinner et al., 1984; Woollacott et al., 1986), as has its role in increasing the incidence of falls (Lord et al., 1999; Overstall et al., 1977; Sorock and Labiner, 1992; Tinetti et al., 1988). The underlying theme is that the loss of proprioception is prevalent with old age in the ankle, knee, and upper-extremity joints. The mechanism through which these changes occur are both central and peripheral. Central mechanisms involve feedback loops between sensory and motor areas (McCloskey, 1978), whereas the peripheral mechanism includes alterations in cutaneous, articular, and Golgi tendon organ receptors. Each of these mechanisms have been investigated in various papers (Aydog et al., 2006; Iwasaki et al., 2003; Morisawa, 1998; Ribeiro and Oliveira, 2007).
Aging-induced reductions in muscle strength and function begin in the second decade. The changes in gait with aging have largely been attributed to muscle weakness. Muscle weakness in turn can be due to alterations in the muscle structure and motor unit complex within one or more muscle groups. The studies in this domain can be grouped into three broad categories depending on their focus: muscle strength, muscle structure physiology, and neuromuscular physiology.
Changes in gluteal muscles strength have been observed in healthy women, with hip extensors showing a decrease from the fourth decade onward and hip abductors showing a decrease from the fifth decade onward (Akbari and Mousavikhatir, 2012). Hip flexor and extensor maximal isometric strength has been reported to be reduced to about 22% and 31% in older women as compared to young women (Dean et al., 2004). The changes in maximal isometric voluntary contractions, especially at the ankle, have been shown to be affected more by changes in muscle fibers and motor unit physiology (Vandervoort and McComas, 1986). Although this provides some insight into strength changes, these results cannot be directly applied to incidences of falls when older adults are undergoing dynamic transformations in posture. In addition, gait does not require lower-extremity muscles to produce forces at their maximal capacity. While climbing stairs, the ankle joint power exerted by both young and old adults was found to be similar. However, when ankle power was normalized to the maximal isokinetic power at the closest ankle angular velocity of a stair gait, it was found that older adults were operating close to their maximal capacity. Therefore, any additional stress occurring as a result of perturbation or loss of balance will increase the risk of falls because muscles may not be able to provide extra power to resist the loss of balance. This provides a direct clinical evidence that there is a loss of functional strength in older adults and that absolute measure of strength does not necessarily provide a complete picture of an individual’s performance or susceptibility to adverse events.
The preceding findings are supported by evidence of slower contractile properties and slower action potential discharge rates in the muscles of older adults. Muscles of older adults have slower contractile properties, which has been demonstrated by smaller rates of maximal torque development and decay within the tibialis anterior muscle of older adults as compared to young (Baudry et al., 2005). This occurs in conjunction with lower rates of action potential discharge and smaller torque generation in ankle dorsiflexors during rapid submaximal contractions in older adults (Klass et al., 2005, 2007, 2008). Two situations arise out of this evidence: (1) if this phenomenon can be extrapolated to other muscle groups critical in maintaining the gait velocity and balance and (2) if this is restricted to ankle dorsiflexors alone or affects ankle dorsiflexors more than any other muscle group associated with continuance of the gait. The first scenario clearly provides an explanation for an increased risk of fall both as a result of failure to carry the movement (because of velocity) and to maintain balance. The second scenario provides a potential reason for a shift in balance strategy in older adults from ankle dominant to hip dominant, especially when encountering a challenging environment or condition such as a ramp (Casebolt and Singhal, in preparation).
Muscle structural changes can occur because of sarcopenia or atrophy. Sarcopenia—or loss of muscle mass and its contractility occurring as a result of old age—results from changes in tissue matrix and cellular structure, which are distinct from the changes that occur due to chronic lack of use, which leads to atrophy. These changes are beyond the scope of this chapter, but both sarcopenia and muscle atrophy cause reduced muscle strength and increased susceptibility to muscle injury. Contraction-induced damage and subsequent regenerative capacity of type II muscle has been reported to be most affected in old adults (Faulkner et al., 1995; Schultz and Lipton, 1982; Singh et al., 1999). It has also been shown that the internal environment of the muscle rather than the myofibrils themselves have an effect on muscle regeneration (Carlson and Faulkner, 1989; Conboy et al., 2005) such that new myofibrils are thinner and more fragile, which may result in increased susceptibility to contraction-induced injury (Renault et al., 2000). Another study has observed that there is a greater amount of atrophy in type II fibers and not a loss of number of fibers with old age (Klein et al., 2003). This does not mean that type I fibers do not undergo changes because cross-sectional areas for both type I and type II vastus lateralis fibers have been found to be decreased in older women (Hunter et al., 1999). These studies highlight not only that the capacity to generate quick and high forces is reduced in older adults but also that their susceptibility to and recovery from an injury increases in cases of sudden perturbations.
The role of strength training in improving muscle strength has been well documented. The guidelines for dosage published by the American College of Sports Medicine can be used to provide a starting point, but an older adult needs to be properly evaluated in order to rule out any other underlying pathology that may require alterations to any rehabilitation plan. Besides an obvious increase in muscle strength, high to moderate resistance training has been associated with increases in antiinflammatory IL-1ra in males and the reversal in pretraining expression of 179 genes (Forti et al., 2016; Melov et al., 2007). These gene expression changes were similar to characteristics of a younger population, implying that resistance training not only reduces inflammatory markers but also reverses certain aspects of aging. A systematic review compared functions in older adults after strength training or power training and found that power training results in better functional outcomes, but the review could not come to any conclusion regarding the safety of any of the methods (Tschopp et al., 2011).
Changes in muscle structure are accompanied by changes in motor unit physiology. A motor unit comprises a motor neuron and the muscle fibers it innervates. All muscle fibers in a motor unit are of one type, and all show contractions at the same time once there is a stimulus. During old age, the changes in motor unit occur in the number of motor neurons available (Campbell et al., 1973), which will consequentially result in some muscle denervation. Some of these denervated fibers undergo degeneration, but there is evidence that motor units are redistributed, with some of these fibers being reinnervated by other motor neurons (Payne et al., 2006). Thus, there is an overall loss in the number of motor units and an increase in twitch force production from the redistributed motor units. However, these changes affect fine motor activity in hand muscles more than the larger muscle groups utilized in gait. On the lower-extremity front, plantarflexor force production abilities have been shown to be associated with walking velocity (Clark et al., 2013). The rate of force production and electromyography (EMG) muscle activation during rapid maximal heel raise is associated with the fastest walking velocities in older adults. The adults who had a lower fastest velocity showed much lower rate of force and EMG activation in medial gastrocnemius. The muscle cross-section area in medial gastrocnemius was similar in both the groups. Although the study was able to show independent associations between walking velocity and ankle plantarflexors force generation capabilities, these are not similar to dynamic walking tasks and one has to be cautious while inferring that limited plantarflexor neuromuscular activation and force generation limits walking speed.
Age-associated changes have been observed in the soleus H-reflex (the EMG equivalent of the stretch reflex) in different standing positions and reduced modulation in reciprocal inhibition of muscles (between agonist and antagonists) during low-force contractions (Kido et al., 2004; Tsuruike et al., 2003). However, changes in the H-reflex were not seen during walking or in complex standing tasks (Chalmers and Knutzen, 2002; Kido et al., 2004; Mynark and Koceja, 2002). These findings again reinforce that the human gait is intrinsically a very different process than standing or any other postural task. Therefore, findings obtained during other activities should be cautiously applied when drawing similarities with gait. Discussion on the presence of spinal circuitry that may aid in controlling locomotion (also referred to as the central pattern generator) is beyond the scope of this chapter.
Imaging studies have shown increased activation in older adults’ cortical and subcortical areas while doing simple upper-extremity motor tasks. Similar cortical adaptations may be associated with observed changes in voluntary activation of the quadriceps femoris during isometric contractions (Klass et al., 2007, 2008). Here again we need to be careful in interpreting the relationships to actual walking tasks. As far as walking or stepping is concerned, evidence suggests the presence of coactivation of agonist and antagonist, which may reduce the overall strength of contraction (Hortobagyi and DeVita, 2000). This coactivation is controlled by descending motor pathways independently of agonist activation (Hortobagyi and Devita, 2006; Levenez et al., 2005), which suggests an adaptation in higher cortical control. The presence of coactivation of muscles works to reduce the net force of contraction acting at the joint, which provides an additional mechanism of strength loss. In addition, coactivation of muscles at a joint is an important mechanism for increasing joint stability, so it may be correlated with changes in proprioception. However, more research is required to establish direct correlation in CNS activity and gait changes.
In conclusion, the underlying physiological changes occurring among the elderly begin as early as the fourth decade but do not become recognizable until much later in life. As a result, the elderly are more likely to slow their gait velocity to accommodate changes in muscle strength, aerobic capacity, cognitive function, and overall physical health. Gait adaptations in old age can be due to age or any other underlying pathology. It is extremely difficult but important to differentiate the cause of these changes clinically in order to ensure an older individual’s proper rehabilitation. A decrease in gait speed increases the chances of a fall. As a person ages, there seems to be a transition from an ankle strategy to one that places increased emphasis on the hip, which is primarily responsible for carrying the upper body during ambulation; therefore, utilizing a hip strategy to produce forward progression while balancing the upper body may prove too demanding if a perturbation should occur. These changes in gait occur due to the conjunction of altered physiology and body structure. Altered sensations affect balance, deteriorating neuromuscular processes affect muscle strength characteristics, and changes in CNS affect overall control and execution. Numerous studies have been conducted to elucidate these effects, but care should be taken to infer whether these can be applied to walking tasks. The studies correlating changes in CNS and locomotion are limited both in number and scope due to the difficulty of measuring CNS activity while a person is ambulating. Technological advances, especially in ambulatory electroencephalography and near-infrared spectroscopy, have made this a possibility, and significant research in this arena can be expected in the near future.