Cycling has advantages in reducing the knee joint loads compared to walking (D'Lima et al., 2008; Kutzner et al., 2012), thus it is frequently prescribed as a rehabilitation exercise by many health professionals (Naal et al., 2007; Salacinski et al., 2012; Walker et al., 2015). Despite the relatively lower joint load during cycling, the prevalence of chronic cycling injuries can be as high as 85% due to poor preparation, equipment and technique as well as overuse (Dettori and Norvell, 2006; Wanich et al., 2007). Among the joints of the lower limb, the knee is thought to be the most affected site with injury prevalence of 42%-65% (Conti-Wyneken, 1999; Dannenberg et al., 1996; Wilber et al., 1995).
The internal knee abduction moment (KAM) is a surrogate measure for loading to the medial compartment of the knee joint and has been shown to be closely associated with the development of medial knee osteoarthritis (OA) in walking (Mundermann et al., 2004; Sharma et al., 1998; Zhao et al., 2007). Studies have shown that the frontal plane knee malalignment can significantly affect KAM during walking in healthy populations (Barrios et al., 2009; Stief et al., 2011) and knee Oa patients (Messier et al., 2014; Turcot et al., 2013). Furthermore, longitudinal studies have shown that varus and valgus alignment were associated with incident and progression of medial and lateral knee OA (Felson et al., 2013; Sharma et al., 2010; Sharma et al., 2001).
Although many studies have investigated the effects of knee alignment during walking, there are a few that have used cycling. Recently, Gardner et al. (2015) compared the KAM in patients with medial knee OA and healthy controls during stationary cycling and found no significant difference between groups. However, knee alignment of the participants was not measured in the study, and it is possible that the knee alignment data may help explain their results on KAM.
Many stationary bikes in fitness or physical therapy facilities have toe clips available and they are used to constrict feet on the pedals during cycling. However, previous studies have suggested allowing some freedom between the foot and pedal may be beneficial for reducing overuse knee injuries (Boyd et al., 1997). It is still unclear whether a toe clip would have any negative effects on knee biomechanics and its effect in individuals with different knee alignment during stationary cycling.
Despite the importance of knee alignment, no studies have explored its effect on knee frontal plane biomechanics during cycling. It is reasonable to assume that the knee alignment may have a similar influence on knee biomechanics during cycling. Furthermore, it remains unclear whether using a toe clip would negatively influence the frontal plane loading in the knee joints. Therefore, the purpose of this study was to examine the effects of knee alignment and use of a toe clip on the knee frontal plane biomechanics during stationary cycling. It was hypothesized that 1) participants with a varus alignment will have a greater KAM compared to participants with a neutral or valgus alignment, and 2) KAM will not differ regardless usage of a toe clip.
Thirty-two participants (11 varus alignment, 11 neutral alignment, and 10 valgus alignment) participated in the study (Table 1). Participants were recruited from a campus student population through flyers and announcements that were made in Kinesiology classes and in Physical Education and Activity Program classes. Using an effect size of 1.091 calculated from the external knee adduction moments in the study by Barrios et al. (Barrios et al., 2009), a total sample size of 10 (or 5 per group) was estimated with a P level of 0.80 and [alpha] level of 0.05 (3.1.3, G*Power) (Faul et al., 2007). An anteroposterior full limb radiograph was obtained to measure the knee mechanical axis angle (MAA). Neutral, valgus, and varus groups were determined as 180[degrees] [+ or -] 2, > 182[degrees], and
All potential participants attended a full-limb radiographic measurement session. The anteroposterior view of a full-length lower extremity weight-bearing radiograph was captured with the graduated-grid x-ray cassette (Moreland et al., 1987; Sharma et al., 2001). The cassette size was 130.0 cm (height) by 36.0 cm (width). The participant stood barefoot with knees in full extension and the tibial tubercles facing the x-ray beam. The x-ray tube was placed at a distance of 1.83 m from the cassette. The x-ray power settings were 95 kilovolts and 300 mA/s-500 mA/s, depending on the limb size and tissue characteristics.
A nine-camera motion analysis system (240 Hz, Vicon Motion Analysis Inc., UK) was used to obtain three dimensional (3D) kinematics during the test. Reflective anatomical markers were placed bilaterally on the acromion processes, iliac crests, anterior superior iliac spine, posterior superior iliac spines, greater trochanters, medial and lateral epicondyles, medial and lateral malleoli, 1st and 5th metatarsal heads, tip of the second toe, and midpoint of the front edge of both pedals. A cluster of four tracking markers on a thermoplastic shell was attached to the shanks, thighs, pelvis and trunk. Three discrete markers were also attached to the lateral, superior and inferior heel counters of the standard lab shoes (Noveto, Adidas). Three lateral pedal markers were also used as tracking markers for both pedals (Figure 1) (Fang et al., 2016). A crank tracking marker was placed on the axes of both crank arms, and an additional tracking marker was placed on the front body of the bike (Fang et al., 2016; Gardner et al., 2015).
A mechanically-braked Monark cycle ergometer (818E, Monark, Sweden) was used in this study. The saddle height on the bike was set so that the angle of the knee joint was approximately 30[degrees] when the crank was at bottom dead center (Bini et al., 2011). The anterior-posterior position of saddle was set so that the knee was in line with the pedal spindle by a meter stick when the crank was in the forward horizontal position (Burke, 2003). The position of the handlebars was set so that the angle between the participant's trunk and thigh was 90 [degrees] when the crank was in the forward horizontal position (Fang et al., 2016; Gardner et al., 2016).
A customized instrumented bike pedal was used on the cycle ergometer, which allows recordings of three dimensional pedal reaction forces (PRF) and moments (Fang et al., 2016; Gardner et al., 2015; Martin and Brown, 2009). The assembly contained two 3D force sensors (Type 9027C, Kistler, Switzerland) coupled with two industrial charge amplifiers (Type 5073A, Kistler, Switzerland). The charge amplifiers were necessary to convert the charge measured by the force sensors to a voltage value used by the Vicon Nexus software. The sensors were placed in the left pedal and a dummy pedal of the same mass and design was used on the right side to minimize asymmetries during the testing. Pedal analog data along with marker data were exported to Visual3D for further analyses. To align the pedal coordinate system with the lab coordinate system, a customized jig was used to secure the front support base of the bike onto a force platform so that the y-axis of the pedal coordinate system was set parallel to the y-axis of the global coordinate system (Figure 1)(Fang et al., 2016).
Participants performed a 2-min warm-up on the cycle ergometer followed by a 2-min rest period. Participants then cycled for 2-min in each of six cycling conditions with a 2-min break between conditions: pedaling at a cadence of 80 rpm and workloads of 0.5 kg (40 Watts), 1.0 kg (78 Watts), and 1.5 kg (117 Watts) with and without a toe clip. All the conditions were randomized by toe clip conditions first, and followed by workload conditions. Simultaneous recordings of kinematic (240 Hz) and kinetic (1200 Hz) data were performed on five consecutive pedal cycles which began during the last 30 seconds of each test condition (Fang et al., 2016; Gardner et al., 2015).
Data and statistical analyses
The obtained radiographs were analyzed using InteleViewer software (Intelerad, Montreal, Quebec, Canada). A 2.54 cm diameter sphere was used to calibrate radiographs. The mechanical axis of each limb was then determined using the following standard procedures (Moreland et al., 1987). The mechanical axis of the femur was measured by a line drawn from the center of the femoral head to the center of the tibial intercondylar eminence and the mechanical axis of the tibia was from the center of the intercondylar eminence to the center of the talus (Bennett et al., 2017a; 2017b). The mechanical axis angle of the knee joint was measured by the medial angle between the mechanical axes of femur and tibia. Two investigators (HB and GS) independently performed the same measurements on each radiograph. Inter-rater reliability, as measured by intra-class correlation (ICC), showed that the average ICC was 0.998 with a 95% confidence interval of 0.996-0.999 (F(33,33) = 4982.280, p
Pedal reaction forces (PRF), moments of force, and center of pressure (COP) on the left pedal were computed in Visual 3D (C-Motion Inc.). Pedal reaction forces (PRF) were computed as the respective sums of the vertical (Z), anteroposterior (Y) and mediolateral (X) components of the 3D force sensors. The moments were computed based the PRFs and distance measurements between the two force sensors and of the pedal. The mediolateral center of pressure (COP) was computed based on the equations provided by the manufacturer. The anteroposterior COP displacement was assumed to be fixed along the pedal spindle. The computed PRFs, moments and COP were then transformed into the lab coordinate system for inverse dynamics calculations. The hip joint center was estimated using the Bell method (Bell et al., 1989). A right-hand rule was used to determine the...