Lowering intraocular pressure (IOP) remains the only proven method of preventing the onset and progression of glaucoma, yet the role of IOP in the disease remains controversial. This largely arises from the wide spectrum of individual susceptibility to IOP wherein a significant number of patients with normal IOPs develop glaucoma (e.g., normotensive glaucoma), and other individuals with elevated IOP show no signs of the disease. It is therefore important to understand the relationship between IOP and glaucomatous optic neuropathy when IOP is only one of several factors that influence the disease ( Box 20.1 ). IOP is, by definition, a mechanical entity – the force per unit area exerted by the intraocular fluids on the tissues that contain them. Glaucomatous optic neuropathy is a biologic effect – likely the result of an IOP-related cascade of cellular events that culminate in damage to the retinal ganglion cell (RGC) axons. One of the challenges of biomechanics is to understand how the mechanics are transduced into a biological response and/or tissue damage. How does this take place? Why is there such a wide range in susceptibility to IOP? Why does elevated IOP lead to that particular cascade of events and not another? What can we do to predict, detect, and stop the progression of glaucomatous damage? Unfortunately we do not have answers to these questions, but in recent years there has been considerable progress towards understanding of the role of IOP in glaucoma. In this chapter, we focus on two main themes: what is known about how IOP-related forces and deformations are distributed in the posterior pole and optic nerve head (ONH), and what is known about the response of the living system.
Intraocular pressure (IOP) is the principal risk factor for the onset and progression of glaucoma
IOP is a mechanical entity that induces glaucomatous damage through a variety of mechanisms
Biomechanical analysis of the effects of IOP on the optic nerve head and sclera will help us understand how IOP leads to glaucoma, and why some individuals are more susceptible to IOP-induced glaucomatous damage than others
From a biomechanical perspective, the ONH is a natural site of interest when studying IOP effects because it is a discontinuity in the corneoscleral shell. Such discontinuities are often weak spots in mechanically loaded systems because they give rise to significant stress concentrations. In addition, it is the ONH, the lamina cribrosa (LC) in particular, that is the principal site of RGC axonal insult in glaucoma. Nevertheless there is evidence both for and against direct, IOP-induced damage to the retinal photoreceptors, and it is likely that there are also important pathophysiologies within the lateral geniculate and visual cortex.
We have proposed that the ONH be understood as a biomechanical structure, and that the mechanical effects of IOP on the tissues of the ONH, namely forces and deformations, are central determinants of both the physiology and pathophysiology of the ONH tissues and their blood supply at all levels of IOP ( Figure 20.1 ). Within this framework, the susceptibility of a particular patient’s ONH to IOP-related insult is a function of the biomechanical response of the constituent tissues and the resulting mechanical, ischemic, and cellular events driven by that response. Experienced over a lifetime at physiologic levels of IOP, these events underlie normal ONH aging. Hence, eyes with a particular combination of tissue geometry and stiffness may be susceptible to glaucomatous damage at normal IOP, while others may have a combination of ONH tissue geometry and stiffness that render them impervious to any deleterious effects of high IOP.
We believe that the mechanical and vascular mechanisms of glaucomatous injury are inseparably intertwined: IOP-related mechanics determines the biomechanical environment within the ONH, mediating blood flow and cellular responses through various pathways. Reciprocally, the biomechanics depend on tissue anatomy and composition, which are subject to change through cellular activities such as remodeling. The interrelationship between mechanics and physiology could be particularly strong within the LC due to its complexity. The LC is composed of a three-dimensional (3D) network of beams of connective tissue, many containing capillaries, that provides functional and structural support to the RGC axons. IOP-related forces within the LC could deform the beams containing capillaries, diminishing the blood supply to the laminar segments of the RGC axons. Conversely, primary insufficiency in the blood supply to the laminar region could introduce cell-mediated connective tissue changes that could remodel the extracellular matrix (ECM), making the laminar beams more prone to failure, and limit the diffusion of nutrients to adjacent RGC axons.
Before proceeding with a presentation of the mechanical effects of IOP on the eye and ONH, we review the basic concepts of mechanics relevant to this analysis.
Basic concepts in mechanics
The following are fundamental terms and concepts from mechanics that may not be familiar to clinicians ( Box 20.2 ). The interested reader may pursue these ideas in greater depth by referring to appropriate textbooks.
Strain is a measure of the local deformation
Stress is a measure of the forces per unit area
Stress and strain are related to each other through the material properties
Ocular tissues are complex materials that are nonlinear, anisotropic, viscoelastic, and inhomogeneous
A nonlinear material varies in stiffness as it deforms
An anisotropic material exhibits different stiffness in different directions
A viscoelastic material exhibits higher stiffness when loaded quickly rather than slowly
Stress is a measure of the forces transmitted through, or carried by, a material or tissue. Specifically, stress is the force divided by the cross-sectional area over which it acts. For example, pressure is a stress and can be expressed in pounds per square inch (psi).
Strain is a measure of the local deformation induced by an applied stress. It is computed as the change in length of a material divided by its resting length, and is often expressed as a percentage. For example, a wire that was originally 10 mm long that is stretched an additional 1 mm exhibits 10% tensile strain.
In addition to tension and compression, a material can undergo shear. Tension, compression, and shear are often referred to as the three modes of stress and strain. However, these three modes are not independent, as shown in Figure 20.2 .
It has been established that the biologic response of tissues and cells depends strongly on the mode of the strain stimulus (tension, compression, or shear), as well as on their magnitudes and temporal profiles. It is therefore of interest to determine which modes of strain and stress the tissues of the ONH are exposed to as IOP is elevated. Note that strains are generally not homogeneous. When the LC deforms, some regions could be highly strained in different modes, whereas others remain largely unaffected. This is important because the biological effects on cells are likely to be strongly dependent on the local levels of strain or stress rather than on global levels.
We would also like to emphasize that mechanical stress, which represents forces, is not synonymous with notions of stress typically used in physiologic or metabolic contexts (e.g., ischemic or oxidative stress). Mechanical stress cannot be measured directly, and we believe it is strain that damages tissues. However, stress is often used to predict the sites of and failure in engineering structures and has been correlated to damage in tissues, so it may be that, while strain is causing the damage, stress is a better predictor of the sites of that damage.
Stress and strain (i.e., forces and deformation) in a material are related to each other through material properties, and this constitutive relationship is intrinsic to each material. For a given load, a material that exhibits large strains is thought of as compliant. Conversely, a material that exhibits small strains for the same load is termed stiff. A stiff tissue such as sclera can have high stress but low strain, while an equal volume of compliant tissue, like retina, might have high strain even at low levels of stress. The simple description above does not account for many of the complexities in material properties that occur in soft biologic tissues, such as anisotropy, nonlinearity, and viscoelasticity ( Box 20.3 ). These complexities are likely to be fundamental to understanding ocular mechanics and will be discussed in the context of scleral biomechanics below.
ONH biomechanics are complex, with the tissues simultaneously subject to tension, compression, and shear
At present, it is difficult to measure the response of the lamina to changes in intraocular pressure (IOP), and therefore computational modeling is being used
These models suggest that IOP-related stresses and strains within the connective tissues of the ONH are substantial, even at normal IOP
From a mechanical perspective, the eye is a pressure vessel, on which IOP produces deformation, strain, and stress. Through computational modeling, the sclera has been shown to have a strong influence on how the LC deforms when IOP changes ( Box 20.4 ). Therefore, understanding the mechanical behavior of the sclera is essential to understanding IOP-induced LC deformations. One of the mechanisms by which scleral biomechanics can influence the response of the LC to IOP is illustrated in Figure 20.3 .
The sclera is the main load-bearing tissue of the eye and deformations of the sclera are transmitted to the optic nerve head at the scleral canal wall
As such, the mechanical properties and behavior of the sclera have a strong influence on how the optic nerve head responds to changes in intraocular pressure
Alas, describing the mechanical behavior of soft tissue such as the sclera is a formidable undertaking that demands extensive experimental and mathematical efforts. The first step in characterizing scleral material properties is the development of an experimental mechanical test to measure the deformation of the tissue subjected to loads (e.g., uniaxial, biaxial, or pressurization tests). The second step is the development of a constitutive model (i.e., relationship between stresses and strains) which describes the tissue mechanical behavior as observed in the experiment and provides a mathematical representation of the tissue’s material properties.
Historically, the sclera has been described as a thin-walled, spherical pressure vessel obeying the analytical equation known as Laplace’s law. Laplace’s law is useful to estimate the state of stress in nonbiological pressure vessels, but it is inadequate for describing many aspects of the eye’s mechanical response to variations in IOP. The sclera has been shown to exhibit several properties that violate the assumptions of Laplace’s law. First, the eye is a pressure vessel of nonuniform thickness. Second, in terms of its material properties, the sclera is nonlinear, anisotropic, and viscoelastic. These concepts are fundamental to understanding scleral and ONH biomechanics, and below we present them in greater detail.
Nonlinearity is a property exhibited by most soft tissues, often as a consequence of collagen fibers within the tissue ( Figure 20.4 ). In a pressure vessel with linear material properties, the stiffness would remain constant during pressure-induced deformation, whereas a nonlinear pressure vessel would experience either softening or stiffening as it deforms. Recent experiments have shown that the sclera exhibits a considerable increase in stiffness, at least fivefold, when exposed to an acute elevation of IOP from 5 to 45 mmHg. This dramatic change shows the substantial impact that collagen fibers can have on scleral stiffness and deformation.
Anisotropy, as opposed to isotropy, is the property by which materials exhibit different stiffness in different directions. For thin biological tissues, anisotropy is primarily dictated by the organization of their fibrous structure, which is confined within the plane of the tissue, as illustrated in Figure 20.4 . Unlike the cornea, scleral collagen fiber orientation has not been fully characterized. However, it has been shown through computational modeling that collagen fiber orientation and distribution are major determinants of scleral deformation. Further experimental work is needed to characterize better scleral anisotropy and its effects on LC biomechanics.
Viscoelastic materials, such as the sclera, exhibit higher resistance to deformation when loaded quickly rather than slowly. Downs and coworkers characterized the viscoelastic material properties of normal rabbit and monkey peripapillary sclera and found that the material properties of peripapillary sclera are highly time-dependent (viscoelastic). This behavior protects ocular tissues from large deformations during short-term spikes in IOP, which occur during blinks, eye rubbing, or high-speed impacts.
These three aspects of the sclera’s material properties – nonlinearity, anisotropy, and viscoelasticity – affect IOP-induced scleral deformations, but they are not the only important determinants of the eye’s mechanical response. Material properties may be combined with thickness and shape to define another useful concept, that of structural, or effective, stiffness.
Studies of the scleral thickness of human and monkey eyes show that, on average, the human sclera is about twice as thick as the monkey sclera. The sclera is thinnest near the equator (as thin as 100 µm in both species) and thickest in the peripapillary region (average of 1000 µm in the human and 450 µm in the monkey). Large variations in peripapillary scleral thickness occur naturally and in pathologic conditions (e.g., myopia ), and have been hypothesized to be an important determinant of individual susceptibility to IOP. Figure 20.5 illustrates how IOP-related stress is distributed in the peripapillary sclera in two situations: homogeneous thickness with a circular scleral canal, and inhomogeneous thickness with an elliptical scleral canal.
Optic nerve head and lamina cribosa biomechanics
Models of the optic nerve head
Initial experimental studies of ONH biomechanics were often designed to examine and quantify a posterior deformation of the LC in response to an acute increase in IOP. Unfortunately, it is difficult to take measurements of the LC directly because it is fragile and relatively inaccessible, and therefore experimental efforts used one of two approaches: histological examination of ONH tissues fixed at different IOPs or measurement of acute deformations of the ONH surface through imaging, and using these deformations as a surrogate for the deformations of the underlying LC.
Both approaches produced interesting results. For example, histology-based studies found that acutely elevating IOP produced a posterior deformation of the LC (12–79 µm in humans, and 10–23 µm in monkeys ). However, these studies also highlighted the large variability in ONH geometry between individuals and the difficulties inherent in histomorphometry, which complicates distinguishing the effects of IOP from the natural differences between eyes. Imaging-based studies were subject to the assumption that IOP-induced deformations of the ONH surface are a good surrogate for deformations of the LC. Some experiments, as well as the modeling studies we describe below, have suggested that IOP-induced deformations of the ONH surface are not good predictors of LC deformations ( Box 20.5 ).
Models have suggested that the intraocular pressure-induced displacements of the optic nerve head surface might not be a good surrogate for those of the underlying lamina
Advances in imaging such as deep-scanning ocular coherence tomography show promise for imaging the acute deformations of the lamina
Recently, there have been advances in direct imaging of the acute deformations of the LC itself using second harmonic imaging or deep-scanning ocular coherence tomography ( Figure 20.6 ). Although promising, these technologies are still in development and have been unable to characterize the response of the ONH tissues to IOP.