The Biomechanical Component of the Subluxation Complex      

This section is compiled by Frank M. Painter, D.C.
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Basic Science Research Related to Chiropractic Spinal Adjusting:
The State of the Art and Recommendations Revisited

FROM:   J Manipulative Physiol Ther. 2006 (Nov); 29 (9): 726–761


Biomechanics is the study of the effects of loads applied to biologic cells, tissues or systems. Biomechanics has its origins from Galileo's studies of mechanics in general and his creation of the term mechanics as a subtitle of his book "Two New Sciences" (1638) to refer to force, displacement, and strength of materials. Arguably, the “father of biomechanics” is Giovanni Alfonso Borelli, who published in "De Motu Animalium" (1681) the principles of muscle movements based on statics and dynamics. However, the word “biology” and its concept as the study of living organisms did not occur until 1802 when the German naturalist, Gottfried Reinhold Treviranus, published his first volume "Biologie; oder die Philosophie der lebenden Natur". To rigorously understand SM and its effects requires an understanding of the principles of biomechanics.

A.   Manipulation Forces

Since the publication of the first white paper in 1997, there have been several important studies that have further clarified the loads that are applied during SM, and especially during high-velocity, low-amplitude (HVLA) SM. Triano and Shultz [150] measured the total force that was transmitted through the body during a side-lying lumbar or lumbosacral HVLA SM. The transmitted forces were similar to the applied forces for their temporal history, but the transmitted forces and moments were shown to vary substantially based on patient positioning. Herzog et al [151] measured the force distribution during thoracic HVLA SM and concluded that there was an important distinction between the total and effective applied forces, with the latter being much smaller than the total applied force. They found that the total peak force was being applied over a mean contact area of 34.8 cm2, but for the thoracic spine, the physiologic contact area of the transverse processes was only 0.25 cm2 (less than 1/100 of the total contact region). Hence, most of the total peak force was being applied to soft tissues (eg, skin, muscle, and fat), and only a small portion (~5 N) was being applied to the transverse process. A similar finding was reported by Kirstukas and Backman, [152] who reported that the “intense contact area” was on the order of 10 cm2 during thoracic HVLA SM. Clearly, the effective applied force during HVLA SM in general will vary based on the contact area of the manipulator's hand and the aspect of the vertebra, but in general, the effective force will be less, and sometimes substantially so, than the applied force.

The 3-dimensional force applied during HVLA SM of the cervical, thoracic, and sacroiliac regions has now been measured. [153] The 3-dimensional data showed that forces in plane with the back (ie, Fx and Fy or shearing forces) always occurred during the SM, which was dominated by the normal (ie, Fz or perpendicular) applied force. The shearing forces were considerable in magnitude, ranging from a low mean of 15% (at T4–5) to a high of 29% (at sacroiliac) of the peak Fz force. As has been previously reported by others, [154] there was a consistent drop in the preload force magnitude just before the impulse portion of the HVLA SM, which is speculated to be due to a “countermovement” affect.

The role of sex in developing force magnitude has been investigated. [155] The only previous report that compared male and female manipulators found no significant differences during HVLA SM using a patient simulator. [156] Forand et al [155] used an experienced matched group (range, 1–24.5 year of experience) of female and male chiropractors (14 per group) and found that there were no significant differences between sexes in thoracic HVLA SM forces. The one exception was that, in the lower thoracic spine, men applied significantly greater preload than did women.

Another type of SM is mobilization or low-velocity low-amplitude SM, which is commonly used by physical therapists as well as other health professionals, including chiropractors. The general approach is to apply an increasing force over 5 to 10 seconds to determine the “end feel,” and then so calibrated, to apply a slow oscillation (~1 Hz for 10 seconds) about a mean graded force (I-IV arbitrary scale), which is less than the end feel. [157] Using an instrumented mobilization table, it was found that there was considerable variation in the force magnitudes used by experienced therapists for end feel, as well as grades I-IV mobilizations of L3 vertebra in healthy subjects. [157] When comparing treatment of younger vs older healthy subjects, it was found that, although mean forces were similar, smaller amplitudes and higher frequency of oscillations were used with older patients. [158] In a study of patients with nonspecific low back pain, there was considerable variation in the magnitudes of forces used, but the variation was strongly influenced not by the patient's severity of complaint but by the physical therapist's training. [159]

B.   Effects of External Loading on Vertebral Displacements

Our understanding of the kinematics of SM has been increased by 2 different types of investigations. First, Keller et al [160–162], have published 2 studies using mechanical force, manually assisted, short-lever SM (ie, Activator [Activator Methods International, Phoenix, Ariz] or very HVLA [VHVLA]) in vivo on patients undergoing lumbar surgery. Using forces ranging from 30 N (lowest setting) to 150 N (maximum setting) on the adjusting instrument, the vertebra where the force was applied had peak displacements of approximately 0.5 mm occurring within 10 milliseconds. Intersegmental displacements occurred of similar magnitudes but with large oscillations lasting 2 to 3 times longer (ie, 20–30 milliseconds), but all oscillations appeared to have damped out within 100 milliseconds. In a second in vivo study, they found that the vertebral displacements due to the Activator instrument were slightly larger (mean, ~0.62 mm) and did not vary significantly, depending on whether the instrument was positioned over the spinous process or facet joint (left or right). [163]

0 Second, using intact cadaveric human lumbar spine specimens, Ianuzzi and Khalsa [164] simulated lumbar HVLA SM while measuring vertebra kinematics and facet joint capsule strain. During simulated HVLA SM, the applied loads were within the range measured during in vivo HVLA SM. Vertebral translations occurred primarily in the direction of the applied load and were similar in magnitude (on order of 1–2 mm) regardless of manipulation site. Vertebral rotations (on order of 1°–3°) and facet joint capsule strain magnitudes (on order of 5%) during simulated HVLA SM were within the range that occurred during physiologic motions. [165] At a given facet joint capsule, distal manipulations induced capsule strains similar in magnitude to those that occurred when the manipulation was applied proximally.

The mobility of lumbar vertebrae in healthy volunteers during mobilization has been assessed using dynamic MRI. Powers et al [166] found that applying a 10–second grade IV posterior to anterior (PA) mobilization (~100 N force) at the spinous process of a lumbar vertebra produced an extension of the vertebra ranging from a mean of 1.2° at L2 to 3.0° at L5. Using plain film radiographs, Lee and Evans [167] found similar displacements for a 150 N PA mobilization at L4. Kulig et al, [168] also using dynamic MRI, found that applying a PA mobilization induced intersegmental motion in all lumbar vertebrae, caudal and cranial, to the site of applied force. This is consistent with the findings of Ianuzzi and Khalsa [164] who also found that simulated HVLA SM at a single vertebra induced motion in all other lumbar vertebrae. Thus, it is not possible to move only a single vertebrae with SM (high or low velocity) because the spine is a linked and coupled structure.

Other effects

HVLA SM is commonly associated with a “cracking” sound, which has previously been shown to be associated with a cavitation phenomenon in the facet joints. [169, 170] In healthy volunteers, Ross et al [171] found that single HVLA SM were typically associated with multiple cavitations (ranging from 2 to 6), which were from nearby vertebrae. This was consistent with the findings of Beffa and Mathews, [172] who found no significant relationship between the location of the cavitation and HVLA SM of the L5 or sacroiliac joint in asymptomatic volunteers. There is some question as to whether HVLA SM can actually induce motion into the sacroiliac joint, as Tullberg et al, [173] using stereo radiography, were unable to measure any significant motion of the sacrum relative to the ilium after a combination of HVLA and mobilization SM in patients with “subluxated” sacroiliac joints. Furthermore, Flynn et al [174] found no association between an “audible pop” and improvement in ROM, pain, or disability in patients with nonradicular low back pain.

C.   Measures of Pathologic States

An intriguing question has begun to be answered relating to whether changes in intersegmental stiffness can be discerned using clinically available tools. Colloca et al [175] measured intersegmental impedance (dynamic stiffness) of lumbar vertebrae and correlated it with characteristics of vertebral height and IVD height measured from plain film radiographs. They found that there was a correlation between decreased disk height at L5–S1 and increased dynamic stiffness at the same segment. These findings were analogous to those of Kaigle et al [176] who, using a porcine model, also observed increased spine dynamic stiffness associated with degenerated disks, compared with normal controls.

Using ultrasound indentation, another noninvasive approach, Kawchuk et al [177] also found that IVD degeneration in a porcine model resulted in decreased indentation for the same applied load. This is an analogous metric as spine stiffness. The use of ultrasound indentation in this animal model had high sensitivity (75.0%), specificity (83.3%), and accuracy (77.1%), compared with other approaches (arthroscopy, MRI, and plain film radiography).

Two biomechanics studies have been performed to examine the effects of fixation (ie, a hypomobile subluxation) of the lumbar spine. Cramer et al [13] used a rat model of fixation in the lumbar spine by externally fixating the spinous processes of L4–L6 for up to 8 weeks. A principal finding due to the fixation was the development of osteophytes and degenerative articular changes of the facet joints within a few weeks. Reversal of some of the degeneration was observed for joints that were fixated for a short term (~1 week), but after 4 weeks, no reversal was observed. Little et al [178] simulated a hypomobile subluxation in intact, cadaveric human lumbar spine specimens by screwing a plate into the left anterior aspect of the L4 and L5 vertebral bodies. During physiologic motions of the fixated spine specimens for flexion, extension, and lateral bending, the motions at L4–5 were significantly decreased, whereas below and above that level, intersegmental motions were significantly increased. Correspondingly, the plane strains of the facet joint capsules were significantly decreased and increased at and above/below the site of fixation, respectively.

Diagnostic tools or outcome measures

The principal biomechanical “tool” still used by most chiropractors is palpation. As such, there has been a continued investigation into factors that change what is felt during palpation. Humans are relatively good at discriminating different magnitudes of stiffness for purely “elastic” materials. [179] However, the human spine responds as a viscoelastic system, in which the speed of force application changes the apparent stiffness. Nicholson et al [180] have shown that the relatively poor ability of clinicians to accurately estimate spine stiffness magnitudes is likely due to a 50% poorer ability to discriminate viscous components of viscoelastic systems. Latimer et al [181] found that therapists used different forces to discern spine stiffness and, hence, had different internal perceptual scales. By training therapists to use a calibrated stiffness instrument, discrimination of PA stiffness in the spine can be done with relatively high interexaminer reliability. [182] Furthermore, objective instruments have been developed that can reliably measure PA spine stiffness. [183] Perhaps, the most important aspect of using palpation to detect subluxations (ie, a “manipulable lesion”) is standardization of training. [184] When examiners are trained in a standardized fashion, they are able to obtain relatively high interexaminer reliability (? = 0.68) for detecting cervical fixations.

Stiffness of the spine is influenced by many factors. If the ribcage is constrained, then the stiffness measured at T12–L4 can be significantly increased. [185] Change in orientation of an applied load to the spinous process can have small yet significant changes in objectively measured stiffness. [186] Furthermore, because the spine is a viscoelastic system, there will be a preconditioning effect when applying loads, such that after preconditioning the spine with standard mobilization SM, there will be no measurable change in stiffness. [187] There has also begun to be a growing appreciation for the natural (and normal) variability in spine stiffness as assessed by standard ROM tests during a physical examination. Christensen and Nilsson [87] found that in asymptomatic volunteers during a 3-week period, there was an intrinsic variability in ROM of the cervical spine of ± 20°, ± 14°, and ± 12° for flexion/extension, lateral bending, and rotation, respectively. In contrast, repositioning the head to the neutral position, which is related to proprioception, is done with relatively high fidelity over the same period. [188] Asymptomatic volunteers were able to reachieve the neutral zero position of their heads with a mean difference of 2.7°, 1.0°, and 0.7° for the sagittal, horizontal, and frontal planes, respectively.

Using a case study approach, Lehman and McGill [189] observed that a single HVLA SM session in the lumbar spine caused notable changes in biomechanical factors associated with a complex task (ie, a golf swing in an experienced golfer who had chronic low back pain). In addition to changes in vertebral kinematics, they observed decreased electromyographic (EMG) responses of the associated lumbar muscles. In a subsequent study, Lehman and McGill [190] found that lumbar HVLA SM in patients with low back pain resulted in variable changes in lumbar ROM and associated muscle EMG. The largest changes were associated with patients with the greatest reported pain. In a review of the available literature. Lehman [191] reported that, currently, the best way to discriminate between normal and low back patient groups was using biomechanical tests that assessed “higher-order kinematics during complex movement tasks.” Simpler end ROM tests had poor predictive ability.

Another commonly performed clinical test is measuring leg lengths, especially in the prone position. Using a special designed table to minimize friction and allow independent loading of each leg, Jansen and Cooperstein [192] determined that the prone leg length test was reliable for detecting non–weight-bearing asymmetry in leg lengths. Nguyen et al [193] found that there was reasonable concordance (? = 0.6) in determining whether a short leg was present using the Activator protocol. Cooperstein et al [194] found that it was possible to detect a leg length difference of 1.9 mm but recommended that only differences of greater than 3.7 mm should have confidence associated with them.

D.   Mathematical and Computational Models

One of the signs of maturity of any field is the ability to produce predictive models. In spine biomechanics, most models are computationally based and either use finite element approaches [195] or optimization with minimization of an objective function. Analytical approaches have also been performed, which include a linear elastic model of a lumbar motion segment. [196] This model successfully predicted loads born by various ligaments under physiologic loads. Solinger [197] created a model that predicted the dynamic response of L2–L3 to impulsive loads on the order of those used in VHVLA SM. Using a lumped parameter approach, Keller and Colloca [198] created an analytical model that predicted the frequency dependent response of the human lumbar spine to PA forces applied to the spinous processes, as is done during low velocity and low amplitude (ie, mobilization), HVLA, and VHVLA (ie, Activator). An alternative approach was adopted by Dulhunty [199] who modeled force transmission in the cervical spine to predict whether parallel forces or concurrent forces are the optimization function. A relatively new approach in spine modeling, especially in the lumbar spine, is to incorporate what are called “follower loads” for muscles. The issue is that the ex vivo (cadaveric) intact lumbar spine will buckle under compressive loads of ~100 N, whereas in vivo, the lumbar spine easily supports compressive loads of ~1000 N (ie, 10 times greater). Patwardhan et al [200] found that by modeling muscle activation so that their loads followed the tangent of the lumbar lordosis, their model would approximate the in vivo condition.

A couple of new comprehensive models have been advanced to explain how the spine becomes subluxated in the first place and how SM can restore it to “normal.” Triano [201, 202] has advocated a mechanical model based on the concept of intersegmental buckling, which was based on original observations by Wilder et al [203, 204] and fluoroscopic recordings of a buckling event in a weightlifter by Cholewicki et al [205] and Cholewicki and McGill. [206] Essentially, this model proposes that there is a balance point between each pair of vertebrae that under certain loading conditions can suddenly shift, which then results in increased tissue strain of associated soft tissues (eg, facet joint capsule). The increased tissue strain can result in small tears and associated biologic inflammatory response. Evans et al [207] have proposed an optimization model where the spine system is biased toward minimizing the mechanical energy associated with loading the spine. Their model is described for the case of linear elasticity, although they claim it is also apropos of nonlinear elasticity. As with any theory (or model), the value of these new theories is really found in their predictive ability and how well their predictions are validated by experimental data. So far, neither of these theories has been tested to any degree.

E.   Instrumented Manipulation

Passive devices have been used for many decades to treat patients with back disorders. Recently, a simple distraction device, Rola Stretcher (Unique Relief, Inc, Davenport, Iowa), designed to be used at home without supervision, was tested to determine whether it showed any lengthening of the spine subsequent to its use. Devocht et al [208] tested 12 asymptomatic adults and found a significant increase in sitting height after 10 minutes of lying supine on the device. They concluded that it at least temporarily lengthened the spine, presumably by increasing the intervertebral disk height.

In addition to the activator adjusting tool, which has had increasing amounts of scientific study, [160–163] the PulStar computer-assisted, differential compliance spinal instrument has been developed, and a few studies on it have appeared. [209, 210] This latter device also applies an impulse load (up to ~150 N), although the duration of the impulse has not been characterized in articles available in the indexed peer-reviewed literature. The device also incorporates a sensor to measure the compliance of the material that it loads, and hence, the compliance of the paraspinal region can be assessed as well as loaded with the same device. A case study has reported that the instrument was used to treat the spines of infants having colic. [210]

F.   Recommendations and Action Steps

  1. Determine (quantify) the biomechanical basis of the subluxation.

    1. Determine the parameters that dictate whether a given vertebra should be manipulated.

    2. Determine the parameters that will guide the optimal approach to administering the manipulation.

  2. Determine the effects of manipulation on tissues of the spine.

    1. Which ligaments (including facet joint capsule) sustain the largest strains due to SM

    2. The influence of the vector direction of a given type of SM on ligament strains

    3. Measure the effects of SM on change in tissue characteristics (eg, ligament modulus of elasticity) and cellular response to SM.

  3. Quantify the biomechanical safety of SM in fracture, disk lesions, ligament strains, muscle, and tendon strains.

  4. Develop comprehensive models of the spine that predict how it responds to physiologic and SM loads.

  5. Determine the biomechanical parameters of SM that dominate the neurophysiologic beneficial effects of SM.

   The Biomechanics of Subluxation   

A Feasibility Study to Assess Vibration and Sound
From Zygapophyseal Joints During Motion Before
and After Spinal Manipulation

J Manipulative Physiol Ther. 2017 (Mar); 40 (3): 187–200 ~ FULL TEXT

Our findings indicated that a larger study is feasible. Other findings included that crepitus prevalence increased with age, was higher in participants with LBP than in healthy participants, and overall decreased after SMT. This study indicated that crepitus assessment using accelerometers has the potential of being an outcome measure or biomarker for assessing spinal joint (facet/zygapophyseal joint) function during movement and the effects of LBP treatments (eg, SMT) on zygapophyseal joint function.

Relationship between Biomechanical Characteristics of Spinal Manipulation and
Neural Responsesin an Animal Model: Effect of Linear Control
of Thrust Displacement versus Force, Thrust Amplitude,
Thrust Duration, and Thrust Rate

Evid Based Compl Alternat Med. 2013 (Jan 20); 492039 ~ FULL TEXT

Like other therapeutic interventions requiring manual deftness, such as surgery, the successful delivery of an HVLA-SM combines knowledge about the motor skills critical for maximizing clinical success and mastery of those motor skills. [7] These motor skills include learning to control the applied force or displacement during the manipulative thrust. With control of these parameters it is important to know whether it is more effective to control force versus displacement and whether there is a thrust amplitude or range of amplitudes critical to producing favorable clinical outcomes? Similarly, is there a specific thrust duration or range of durations over which the peak force or displacement amplitude develops that can make an HVLA-SM most effective? The answers to these questions may be viewed as elements of the manipulation's dosage, conceptually similar to chemical characteristics, such as molecular composition and permeability, which determine a drug's clinically effective dosage.

The Basis for Spinal Manipulation: Chiropractic Perspective
of Indications and Theory

J Electromyography and Kinesiology 2012 (Oct); 22 (5): 632–642

It is reasonable to think that patients responding to spinal manipulation (SM), a mechanically based therapy, would have mechanical derangement of the spine as a critical causal component in the mechanism of their condition. Consequently, SM practitioners routinely assess intervertebral motion, and treat patients on the basis of those assessments. In chiropractic practice, the vertebral subluxation has been the historical raison d'etre for SM. Vertebral subluxation is a biomechanical spine derangement thought to produce clinically significant effects by disturbing neurological function. This paper reviews the putative mechanical features of the subluxation and three theories that form the foundation for much of chiropractic practice. It concludes with discussion of subluxation as an indicator for SM therapy, particularly from the perspective that subluxation may be one contributory cause of ill-health within a "web of causation".

Biomechanical Lesion:
A Better Diagnostic Term for the Profession

Dynamic Chiropractic (September 9, 2012) ~ FULL TEXT

The term spinal biomechanical lesion then would imply a pathological condition involving discontinuity (loss of cohesion) of tissue, and loss of normal vertebral joint function (kinesio-pathology) that often has injury as its cause. It is a very descriptive term, and one that should be well-understood by any and all who have training and study in the field of medical terminology. It's a phrase that is easy to get your head around, no matter your particular discipline.

Biomechanics of Spinal Manipulative Therapy
Spine J. 2001 (Mar); 1 (2): 121–130 ~ FULL TEXT

There currently are a number of named systems of manual procedures. No current triage system is available that predicts which patient has the greater likelihood of benefiting from manual treatment or the procedure type. The biomechanical parameters of SMT form a systematic characterization of manual procedures. Such a system may be used in future studies to test hypotheses of treatment effect from quantitatively defined procedures.

The Functional Spinal Lesion:
An Evidence-Based Model of Subluxation

Topics In Clinical Chiropractic 2001 (Dec); 8 (1): 16–28 ~ FULL TEXT

The buckling model builds on clinical observations and supplements them with both direct and indirect biomechanical evidence. This model does not preconceive or proscribe any source of symptoms, but is able to accommodate the multifaceted clinical presentations of patients who respond favorably to manipulation/adjustment. It also can sustain a variety of hypothetical and evidence-based challenges. These findings offer an opportunity to reconceptualize and refine theoretical models of the spinal lesion into a platform for scientific, clinical, and political advancement of the profession.


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