871
Experimental
31. Experimental Methods in Biological Tissue Testing
Stephen M. Belkoff, Roger C. Haut
31.1 General Precautions.............................. 871
31.2 Connective Tissue Overview ................... 872
31.3 Experimental Methods
on Ligaments and Tendons ................... 873
31.3.1 Measurement
of Cross-Sectional Area ................. 873
31.3.2 Determination of Initial Lengths
and Strain Measurement
Techniques .................................. 873
31.3.3 Gripping Issues in the Mechanical
Testing of Ligaments and Tendons..
31.3.4 Preconditioning
of Ligaments and Tendons ............
31.3.5 Temperature and Hydration Effects
on the Mechanical Properties
of Ligaments and Tendons ............
31.3.6 Rate of Loading
and Viscoelastic Considerations......
31.4 Experimental Methods in the Mechanical
Testing of Articular Cartilage .................
31.4.1 Articular Cartilage.........................
31.4.2 Tensile Testing of Articular Cartilage
31.4.3 Confined Compression Tests ...........
31.4.4 Unconfined Compression Tests .......
31.4.5 Indentation Tests
of Articular Cartilage .....................
874
874
875
875
876
876
876
876
877
877
31.5 Bone ...................................................
31.5.1 Bone Specimen Preparation
and Testing Considerations............
31.5.2 Whole Bone.................................
31.5.3 Constructs ...................................
31.5.4 Testing Surrogates for Bone ...........
31.5.5 Outcome Measures .......................
878
878
879
880
880
880
31.6 Skin Testing .........................................
31.6.1 Background .................................
31.6.2 In Vivo Testing .............................
31.6.3 In Vitro .......................................
883
883
883
883
References .................................................. 884
31.1 General Precautions
Unlike most engineering materials, many biological
tissues are considered biohazards and appropriate precautions must be taken in their handling. Human
cadaveric tissue is considered potentially infectious,
as is nonhuman primate tissue and occasionally other
animal tissue. In the USA, anyone who may be exposed to potentially infectious tissue is required by
Occupational Safety and Health Administration regulation 29 CFR 1910.1030 to receive annual training in
bloodborne pathogens exposure prevention. The training is provided free of charge by employers. Also,
the employer must provide proper personal protective equipment, such as gowns, gloves, facemasks,
etc. Laboratories in which human tissue is used are
Part D 31
The current chapter on testing biological tissue is
intended to serve as an introduction to the field
of experimental biomechanics. The field is broad,
encompassing the investigation of the material
behavior of plant and animal tissue. We have
chosen to focus on experimental methods used
to test human tissue, primarily the connective
tissues ligament, tendon, articular cartilage, bone,
and skin. For each of these tissues, the chapter
presents a brief overview of the structure–function
relationship of the tissue and then discusses some
of the common tests conducted on the tissue to
obtain various material and mechanical properties
of interest. The chapter also highlights some of
the stark differences in testing biological tissues
compared with engineering materials with which
the reader may be more familiar.
872
Part D
Applications
registered as biosafety laboratory level 2 facilities
Centers for Disease Control (CDC). Researchers are
encouraged to check with their local authorities to
be in compliance with laws and regulations governing the use and transportation of cadaveric tissue.
Employers in the USA are also required to offer Hepatitis B vaccinations to those at risk of exposure to
human tissue. Most institutions also require investigators to complete courses on animal handling and
care.
31.2 Connective Tissue Overview
The human body is composed of four primary groups of
tissues:
Part D 31.2
1. Epithelial tissues, characterized by having cells
closely joined one to another and found on free surfaces of the body.
2. Muscle tissues, characterized by a high degree of
contractility of their cells or fibers. Their primary
function is to move the skeleton.
3. Nervous tissues, composed of cells specialized in
the properties of irritability and conductivity.
4. Connective tissues, in which the cells are separated
by large amounts of extracellular materials.
Tissues are combined in the body to form organs.
Organs are defined as structures composed of two or
more tissues integrated in such a manner as to perform certain functions. The heart, for example, has
its chambers lined with a special epithelium, its walls
are primarily muscle, and there are connective tissues
present in numerous forms throughout this organ.
Connective tissues represent wide-ranging types,
both in their variety and distribution. They are all characterized, however, by large amounts of extracellular
material. Their functions are as varied as the tissues
themselves. These tissues bind, support, and protect the
human body and its vital organs. Structural integrity
and function of vital organs are adversely affected when
connective tissue surrounding them is damaged. Connective tissues give us the strength to resist mechanical
forces, and provide a recognizable shape that persists
in the face of these forces. Connective tissues can be
classified by their extracellular constituents. Connective
tissue proper is subdivided into loose, dense, or regular. Loose connective tissue is widely distributed. For
example, it is found in the walls of blood vessels, surrounding muscles as fascia, and in the lung parenchyma.
Dense connective tissues contain essentially the same
elements as loose connective tissue, but there are fewer
cells. This type of tissue is found in skin, many parts
of the urinary ducts, digestive organs, and blood vessels. The stromas and capsules of the internal organs,
e.g., kidney and liver, are dense connective tissues that
are responsible for maintaining the structural integrity
of organs against mechanical forces. Other body components have primarily mechanical functions and are
composed almost exclusively of connective tissue with
few cells called regular connective tissue. Bone tissue
forms the greater part of the skeleton and it primarily
resists the forces of compression resulting from muscle contraction and gravity. Tendons and ligaments have
a parallel arrangement of the extracellular component
that allow these tissues to transmit tensile loads.
All connective tissues are, in fact, complex fiberreinforced composite materials. The mechanical properties of soft connective tissue depend on the properties
and organization of collagen fibers in association with
elastin fibers, which are embedded in a hydrated matrix
of proteoglycans. The constitution of each connective
tissue is tuned to perform a specific function. Ligaments
provide mechanical stability to joints. The function of
tendons is to transmit high tensile forces to bone via
muscles and to allow the muscles to function at their
optimal length. Skin is a two-dimensional soft connective tissue that supports internal organs and protects the
body from abrasions, blunt impact, cutting, and penetration, while at the same time allowing considerable mobility. Articular cartilage is a 1–5 mm layer of connective tissue that covers the articular surfaces of diarthrodial joints. The primary functions of the cartilage are
1. to spread loads across joints and in so doing to minimize contact stresses, and
2. to allow relative movement of the opposing surfaces
with minimum friction and wear.
Bone is the connective tissue that makes up the skeleton of the body. Like other connective tissues, bone
has a significant amount of extracellular material, but
in the case of bone there is a substantial mineral phase
that imparts its characteristic strong and stiff structural
properties.
In biological tissue there is an intimate relationship between structure and function. Some consider
Experimental Methods in Biological Tissue Testing
the structure of a given tissue to have been optimized
through evolution to provide a given function or set
of functions. When subjected to trauma, i. e., impact
or overuse types of injuries, or disease the structure of
the connective tissue is altered or distorted. The altered
structure expresses itself as tissue dysfunction. Therefore, tissue function/dysfunction is often used clinically
as a diagnostic tool for tissue damage. Therefore it is
31.3 Experimental Methods on Ligaments and Tendons
873
important to characterize the mechanical behavior of all
types of connective tissues. The current chapter introduces various experimental methods used to evaluate
the behavior of native and damaged tissue. We will also
introduce experimental methods important in the documentation of tissue subjected to repair processes, either
by natural healing or clinical interventions such as bone
plating, skin suturing, etc.
31.3 Experimental Methods on Ligaments and Tendons
31.3.1 Measurement of Cross-Sectional Area
Accurate measurement of the geometry of ligaments
and tendons is essential in order to determine the material properties of these tissues. In the body, these
structures typically have irregular, complex geometries
that make cross-sectional area measurements difficult.
Generally, the techniques that have been documented
in the literature for determination of cross-sectional
areas of ligament and tendons involve either contact or noncontact methods. Contact methods include
molding techniques, digital vernier calipers, and area
micrometers [31.1, 2]. These methods often rely on
the investigators ability to gently touch the specimen
without causing significant deformations that may expel water and alter the dimension of interest. And
yet, in numerous studies, investigators have also developed contact methods of measurement that rely on
compression of the specimen into a defined shape
by the application of a known, standardized load or
stress [31.3, 4]. One study has developed an area
micrometer technique in which a compressive, external pressure of 0.12 MPa is applied to ligaments and
tendons in the determination of their cross-sectional
area [31.5]. Due to the viscoelastic nature of these tissues, in part resulting from fluid flow within the tissue,
the use of such devices and techniques must incorporate
time as a variable in the measurement of cross-sectional
area.
To minimize the distortion of tissue shape, other
investigators advocate the use of noncontact methods.
These techniques include the shadow technique [31.4],
the profile method [31.6], and the use of a light
source [31.7]. A more recent and often used measurement tool is laser microscopy for the measurement of
cross-sectional area [31.8, 9]. This technique has been
shown to be highly accurate and reproducible. In addition, to correct for errors inherent in the system due to
specimen concavities, a low-cost laser reflectance system has been described [31.10].
31.3.2 Determination of Initial Lengths
and Strain Measurement Techniques
Another difficulty that must be dealt with in the determination of the mechanical properties of ligaments
and tendons is the measurement of specimen length.
Ligaments are attachments between bones, but the insertion points vary over an area. There are direct
insertions through Sharpey’s fibers and there are indirect insertions, in which the ligament fibers merge
with the collagenous tissue of the periosteum. Because
Part D 31.3
Ligaments and tendons are parallel-fibered, dense connective tissues. These complex, fiber-reinforced composite materials provide stability to joints and aid in the
control of joint motion. The fibers, collagen and elastin,
are embedded in various proportions depending on tissue function, in a gelatin-like matrix of macromolecules
(proteoglycans) and water. The role of ligaments, which
connect bone to bone, is to augment the mechanical stability of joints and help control joint function. Tendons,
on the other hand, attach muscle to bone and typically transmit large tensile loads across joints to control
motions of the body. Tendons also allow muscles to
function at their optimal length.
In general, the tensile mechanical response of ligaments and tendons is highly nonlinear and dependent
on the rate of loading or stretch. Such complex mechanical behavior presents a number of challenges when
conducting tissue tests. The following will attempt to
discuss some of the basic concerns and methods needed
in the evaluation of the mechanical (material) and structural properties of ligaments and tendons, based on
previous studies performed with animal and human tissues.
874
Part D
Applications
Video cassette recorder
Video dimensional analyzer
Load
cell
Femur
ACL
Video camera
Tibia
Part D 31.3
Computer
ACL
Tensile load
TV monitor
VDA system
Fig. 31.1 Typical experimental setup of displacement measurement
using video dimension analysis (VDA)
insertions are not discrete, there can be substantial
variance in the initial length of a specimen [31.8,
11]. Various researchers have used pins or wires to
mark the insertions of ligaments into bones. The distances between the markers have been determined
from roentgenographs [31.11–13] or directly with
a ruler [31.14]. Computation of strain can be based
on the deformation of the ligament over these lengths
or grip to grip, however, one should be mindful that
surface strain varies along the length of ligaments and
tendons [31.15]. Local strains can be measured by segmenting the specimen by means of drawing or fixing
reference markers on the surface of the tendon or ligament. The markers can be drawn or fixed on the
surface of a tendon or ligament by means of Verhoeff’s
stain [31.16], elastin stain [31.10, 17], or reflective
tape [31.18]. A charge-coupled device (CCD) camera,
which is part of the video dimension analysis (VDA)
equipment [31.19–21], is then used to record the motion of these markers during tensile stretch and the data
are converted to surface strains Fig. 31.1.
31.3.3 Gripping Issues in the Mechanical
Testing of Ligaments and Tendons
Another reason to use noncontact video systems in
the measurement of ligament or tendon tissue strain
is the potential for specimen slippage within the tissue grips. Actual grip-to-grip strain or surface strain
can be measured more accurately when one of these
systems is employed during testing. When it is possible or appropriate to apply clamps to ligaments or
tendons directly, a number of specially designed freezing, hydraulic, and pneumatic clamps with roughened
gripping surfaces have been utilized. Freeze clamps
have been successfully used in the mechanical testing of
musculo-tendonous junctions [31.22] as well as bovine
and human tendons [31.23]. These types of clamps
maintain a constant pressure against the soft, deforming tissue during axial tensile stretch. In many cases,
however, the substance of a ligament may be too short.
In this case, the entire bone–ligament–bone preparation is utilized for testing. Typically, a normal vice-grip
type of clamp may be sufficient, especially when the
bones can be shaped adequately to fit snugly into standard clamps. In other cases, for example, when testing
a patella–patellar tendon–tibia preparation in which the
bones have a nonuniform shape, the bone end can be
embedded in room-temperature-curing epoxy or bone
cement polymethal methacrylate (PMMA) and inserted
into a holder grip [31.2, 24].
31.3.4 Preconditioning
of Ligaments and Tendons
Biological tissues are viscoelastic and exhibit natural
states in response to repeated application of load or
stretch. Such a state in vivo is called a homeostatic state,
whereas in vitro it is called a preconditioned state. Some
consider preconditioning tissue a necessary step in rheological testing of biological tissues [31.25]. Because
biological tissues are viscoelastic and have memory,
preconditioning specimens means that the specimens
all have the same recent history. If a certain procedure
for testing (stressing or straining) is decided upon, that
procedure should be followed a number of times until the response becomes steady before the mechanical
response of the tissue is documented. If the protocol
changes, for example the amplitude changes, then the
specimen should be preconditioned at that new level.
In most tendons and ligaments, preconditioning effects
on the stress decay during consecutive cycles are assumed to reach a steady response after approximately
10–20 cycles of loading [31.26]. However, a recent
more detailed study of the preconditioning phenomenon
suggests that the effect can persist in some tissues for
many more cycles. It has therefore been suggested that
preconditioning has to be integrated into constitutive
formulations of biological tissues [31.27]. However,
while the nonlinear and viscoelastic aspects of many
Experimental Methods in Biological Tissue Testing
tissues such as ligaments and tendons are well documented in the literature, the effects of preconditioning
or its mechanisms are not well understood [31.28].
31.3.5 Temperature and Hydration Effects
on the Mechanical Properties
of Ligaments and Tendons
31.3.6 Rate of Loading
and Viscoelastic Considerations
Ligaments and tendons are known to be exposed to
varied loads of deformation (loading) during normal
physiological activities and extremely high rates during
traumatic injury [31.32]. Currently, most of the literature has assumed a constant strain rate of 100%/s for
physiological studies and approximately 1000%/s or
above for traumatic injury studies [31.32]. Thus, it is
important to consider strain rate in the experimental
methods used for the study of ligaments and tendons.
The time dependence that does exist in ligaments and
tendons is largely due to the viscoelastic [31.33] or
biphasic [31.34] nature of these tissues. For the description of viscoelastic properties, there are two relevant
quantities of interest: creep and relaxation. Relaxation
relates to the decrease in load in a tissue under repeated
or constant elongation, while creep relates to the increase in elongation under repeated or constant load.
The quasilinear viscoelastic theory (QLV) is the most
widely accepted model of viscoelasticity for ligaments
and tendons [31.21]. In experimental studies relaxation
is the more commonly measured property.
Studies on bone–ligament–bone preparations have
shown that the failure characteristics of these structures
are highly strain rate dependent. These experiments
have indicated that generally high-strain-rate experiments will produce failure of the ligament substance,
while low strain rates more typically produce failure of
bone near the sites of insertion [31.35, 36]. However,
more recent studies with animal models suggest that the
rate sensitivity of the ligament substance itself may have
been overstated in the early experiments [31.32,37] and
the QLV theory takes the form
t
G(t − τ)
σ (t) =
dσ e (ε) dε
dτ ,
dε dτ
(31.1)
0
where G(t) is the reduced relaxation function and ε(t) is
the strain history parameter [31.38]. While G(t) theoretically must be determined under step changes in strain,
an improved method following a finite ramp time has
recently been documented [31.39]. The inherent elastic
function σ e can vary slightly between tissues, but for
tendon [31.40] and ligament [31.38] it takes the form
σ e = A( e Bε − 1) ,
(31.2)
where A and B are constants that are typically determined during a fast constant-strain-rate test using
a variety of least-squares-based fitting routines.
Ligaments, however, probably function in normal
daily activity under repeated low loads, thus they function through creep rather than relaxation. The QLV
theory has also been formulated in creep [31.28]. However, experimental studies have shown that the stress
relaxation response can only be predicted from creep
if collagen fiber recruitment is also accounted for in
this model [31.41]. Finally, recent studies suggest that
these tissues are, indeed, nonlinear so the currently accepted theory needs modification to include nonlinear
viscoelastic characteristics [31.42].
875
Part D 31.3
Environmental conditions, including temperature and
hydration, are important considerations when testing
ligaments and tendons. Testing specimens in air at
room temperature will yield different results than when
immersing them in an aqueous bath of an isotonic
solution where the pH and temperature are closely controlled. Generally, the stiffness of ligaments and tendons
will increase slightly when the bath temperature is decreased [31.28, 29], and there will be a reduction in the
amount of cyclic stress relaxation when these tissues are
immersed in baths at reduced temperatures. Similarly
the rate of cyclic stress relaxation will be significantly
reduced as the concentration of water in the tissue is
reduced [31.30]. A significant increase is noted in the
modulus and strength of human patellar tendons when
tested in a phosphate-buffered saline (PBS) bath versus
when tested using a PBS drip onto its surface [31.2].
The notion that the extent of tissue hydration plays a significant role in the mechanical properties of ligaments
and tendons has also been confirmed in experiments
in which human patellar tendon is stretched at a high
(50%/s) or low (0.5%/s) rate. At high rates of strain
the structural stiffness of these human tendons is significantly higher when immersed in a hypotonic (high
water content in the tendon) solution versus a hypertonic (low water content) solution [31.31]. In contrast,
for a low rate of strain, the structural stiffness is not
dependent on the tonicity of the bath solution. This suggests that the viscous response is related to the water
content of the specimen and not some inherent viscoelastic property of the collagen fibers.
31.3 Experimental Methods on Ligaments and Tendons
876
Part D
Applications
31.4 Experimental Methods in the Mechanical Testing
of Articular Cartilage
31.4.1 Articular Cartilage
Part D 31.4
There are three broad classes of cartilaginous tissues
in the body: hyaline cartilage, elastic cartilage, and fibrocartilage. These tissues are distinguished by their
biochemical composition, their molecular microstructure, and their biomechanical properties and functions.
Hyaline cartilage is normally glossy, smooth, glistening,
and bluish-white in appearance. This tissue covers the
articulating surfaces of long bones and sesamoid bones
within synovial joints, e.g., the surfaces of the tibia, the
femur, and the patella of the knee joint.
Articular cartilage is vital to maintaining normal
joint motion, and its degradation is key to degenerative diseases such as osteoarthritis. Articular cartilage
in freely movable joints, such as the hip and knee, can
withstand very large compressive loads while providing
a smooth, lubricated, load-bearing surface. In order to
understand the mechanical properties of normal articular cartilage and those of degenerated tissue, a number
of methods have been documented in the literature. The
following will attempt to describe some of these experimental methods. The specific choice of test depends
however on the size, shape, and amount of tissue available for study and the objectives of each study.
From these sheets of tissue, dumbbell-shaped or
rectangular specimens are cut with a stamping device.
The tissue slices are then placed in grips which have
the faces lined with fine sandpaper (approximately 1500
grit) [31.47]. The tensile response of specimens oriented
parallel and perpendicular to split lines is exponentially
stiffening, similar to most other soft biological tissues.
These data can be fit with an equation of the form:
σ = A( e Bε − 1) ,
(31.3)
where A and B are found to be significantly higher near
the surface of the tissue and for those test specimens
oriented parallel to the surface split lines [31.48]. As in
most soft connective tissue testing, preconditioning is
often performed before these experiments. In early studies, constant-strain-rate tests were performed [31.49].
More recently, however, ramp-step relaxation testing is
the method of choice. Typically, the tissue slices are
stretched using a moderately rapid ramp in 2% strain
increments to approximately 10–15% strain. Following
each ramp, stress relaxation is invoked. While earlier
studies (i. e. Woo et al. [31.50]) have equilibrated the
tissue for 10–30 min after each step, a recent study suggests that equilibrium requires several hours [31.47].
Equilibrium values of stress and strain are then documented for various layers of the tissue [31.51, 52].
31.4.2 Tensile Testing of Articular Cartilage
31.4.3 Confined Compression Tests
The tensile properties of articular cartilage are extremely relevant to the compressive stiffness of the
tissue [31.43]. When a strip of cartilage is stretched
under a constant rate, the tensile stress–strain curve behavior is nonlinear.
Specimen orientation is important, because articular cartilage is not isotropic. The primary strength and
stiffness directions follow the so-called split or cleavage
lines according to Hultkrantz, which can be observed
by penetrating an India-inked needle into the surface
layer of the tissue [31.44]. These and other studies have
verified that the mapped lines follow the primary orientations of the strength-bearing collagen fibrils in the
tissue.
Since the concentration, content, and organization
of collagen fibrils in articular cartilage varies significantly with depth into the tissue [31.45], tensile tests are
most often conducted on thin layers of tissue cut parallel to the surface with a sledge microtome [31.46] or
a Vibratome (Scott Scientific, Montreal).
The intrinsic compressive properties of cartilage are
usually obtained by the confined compression creep
test. Typically, a cylindrical plug of cartilage and underlying subchondral bone are placed in a rigid cylindrical
chamber, where deformation can occur only in the direction of loading (Fig. 31.2).
Load
Cartilage
Subchondral
bone
Porous
filter
Fig. 31.2 Sectional view of
a confined compression test
fixture
Experimental Methods in Biological Tissue Testing
31.4 Experimental Methods in the Mechanical Testing of Articular Cartilage
n=0
(31.4)
where u(0, t) is the surface displacement, h is the
specimen thickness, and F0 is the applied load. While
a theoretical solution was documented some time ago
for the relaxation test where step changes in displacement are applied to the specimen [31.53], few
investigations have used this method. One reason may
be the extremely high force response experienced in
the test when the initial displacement input is applied
rapidly to the specimen. This typically yields unsatisfactory results in the curve-fitting process.
31.4.4 Unconfined Compression Tests
More typically, relaxation parameters are calculated
from data obtained on cartilage using an unconfined
compression test [31.54]. In this experiment cartilage
discs are removed from the underlying subchondral
bone and placed between two highly polished parallel
plates (Fig. 31.3). A ramp compression is then applied
to a prescribed level of strain (typically less than 20% of
the tissue thickness) and held until an equilibrium load
is reached. The linear biphasic solution for this problem
has been given by Armstrong [31.55] as
(1 − vs )(1 − 2vs )
σ = E s εc 1 +
(1 + vs )
α2 H k ∞
− n 2A
1
h
e
×
,
(1 − vs )2 αn2 − (1 − 2vs )
n=1
(31.5)
where εc is the applied strain, αn are the roots of
the characteristic equation J1 (x) − (1 − vs )x J0 (x)/(1 −
2vs ) = 0, and J0 and J1 are Bessel functions. For the
case of unconfined creep, the theoretical solution has
also been given [31.56].
Load
h
Cartilage
Polished
surfaces
Fig. 31.3 Schematic of an
unconfined compression test
fixture
Unconfined compression studies have documented
a difficulty in using the linear biphasic model of cartilage for the fitting of experimental data [31.54]. The
difficulty appears to arise from the inability of this
model to adequately represent the lateral constraint generated by the collagen fibrils, which lie parallel to the
tissue surface in the top layer. To account for their
stiffening effect in the tissue under unconfined compression, transversely isotropic models of the solid phase
have been proposed [31.57] and later disputed [31.58]
because the models do not account for the tension–
compression nonlinearity of the fibrils [31.59]. Such
studies have led to more recent developments of computational models in which fibril reinforcements are added
to simulate the effect of tension–compression nonlinearity in collagen fibrils [31.60, 61]. These more recently
developed models adequately represent the response of
articular cartilage in unconfined creep and relaxation
compression tests [31.62].
31.4.5 Indentation Tests
of Articular Cartilage
Indentation tests have been used to characterize the
compressive behavior of articular cartilage. A singlephase linear elastic model is often used when modeling
either short-time response or the long-time equilibrium
response of this tissue [31.63, 64]. In the Hayes et al.
study, elastic solutions are given for the indentation of
a rigid, flat or spherical indenter into a layer bonded to
a rigid half-sphere. The solution for a flat indenter is
4Ga
a
P
(31.6)
=
κ
,v ,
ω
(1 − v)
h
where G is the shear modulus, v is the Poisson’s ratio,
and κ is a correction factor that accounts for the finite
layer effect. A nonlinear correction factor is used when
there are deep penetrations into the tissue where nonlinear effects become more important [31.65]. In order to
determine the shear modulus, Poisson’s ratio (v) must
be either assumed a priori [31.66] or determined by
other means [31.57, 67]. In the latter study, the authors
Part D 31.4
This is a uniaxial test. During loading, fluid escapes
only from the top of the specimen through a porous
platen. In the creep test a constant load is applied to the
specimen [31.28]. Analysis of the steady-state stress–
strain response provides the equilibrium compressive
modulus HA , the aggregate modulus of the solid phase
of the tissue. Other model parameters, such as permeability and Poisson’s ratio, are found by curve fitting
the final 30% or so of the creep response using the
solution according to the basic biphasic model for the
cartilage [31.53], given as
2 2
2
∞
e −(n+1/2) π HA kπ/h
F0
u(0, t)
1−2
,
=
h
HA
(n + 1/2)2 π 2
877
878
Part D
Applications
Force
δ
δ(t) = 0.05 mm/s
t
Tidemark
δ(t)
Cartilage
thickness
Articular
cartilage
h
Time
Subchondral bone
Part D 31.5
Fig. 31.4 Typical results obtained from an indentation test to obtain cartilage thickness. The force trace shows the initial indentor
contact with the cartilage surface (sudden resistance with increased
displacement) and contact with subchondral bone (sudden change
in slope)
used two flat indenters with different radii to determine
G and v from the indentation tests on cartilage. In the
former study, Poisson’s ratio was estimated, based on
experimental test results, to be approximately 0.5 instantly to represent the incompressible nature of the
tissue during the initial step of the relaxation test and
v = 0.4 at equilibrium.
Accurate knowledge of the tissue thickness at the
site of indentation is essential for the extraction of
material properties from this test. The methods for
this measurement have included optical [31.68], needle
probe [31.63, 69, 70], and ultrasonic techniques [31.69,
71–73]. In the needle probe method, a small-diameter
probe is displaced into the tissue until the zone of underlying calcified cartilage or subchondral bone are noted
by an abrupt increase in reaction force (Fig. 31.4).
It should be noted here that the shear modulus obtained from such in situ indentation tests is higher than
that obtained from the in vitro unconfined or confined
compression test, possibly as a consequence of indenter
size [31.61].
The indentation creep experiment has been used
extensively to determine the compressive behavior of
articular cartilage [31.68, 69]. The mathematical solution for indentation creep (and relaxation) using
a porous probe on articular cartilage modeled as a linear
biphasic material has been described [31.74]. Early investigations with this model solution typically used a set
of master curves to approximate the intrinsic parameters
(HA , v, and k) of the cartilage. More recently the problem has been simulated computationally and the model
parameters extracted using a least-squares fitting technique. Various experimental studies have since shown
that, for an indentation creep test on articular cartilage
with a porous probe and using the linear biphasic model,
only about the final 30% of creep can be fitted for the
extraction of HA , k, and v for the tissue [31.75, 76].
In some of the more recent models that include a fibril-reinforced network of collagen fibrils,
the actual experimental curves are better fitted for
the extraction of the intrinsic material parameters for
the tissue [31.43, 77]. The same degree of fit can
also be realized for these experimental curves by
the inclusion of a viscoelastic matrix response in the
model [31.72, 78, 79].
31.5 Bone
Bone, like other biological tissue, is not homogeneous
and isotropic and is a hierarchically organized composite material. It is important to have an appreciation of
the microstructural organization of bone in order to understand the simplifying assumptions. A more in-depth
treatment of bone structure and function may be found
elsewhere [31.56]. There are two types of bone: cortical and cancellous. Cortical bone is the dense compact
outer shell of what we casually call bone. Cancellous
bone is a lattice of trabeculae surrounded by marrow,
typically found at the ends of long bones. While the cortex of long bone thins with age, it is cancellous bone
that is most profoundly affected by aging, particularly
in postmenopausal women, and is associated with os-
teoporosis. Osteoporosis is the state of reduced bone
density resulting from the bone-resorbing cells (osteoclasts) outpacing the bone-forming cells (osteoblasts).
31.5.1 Bone Specimen Preparation
and Testing Considerations
Gripping
One of the great challenges in testing biological tissue
is maintaining a firm grip at the tissue–fixture interface. When testing machined coupons, the gripping
issue is typically resolved using standard engineering
material grips. When testing involves whole bones,
the varied geometry of the bone often poses a chal-
Experimental Methods in Biological Tissue Testing
lenge. Gripping can be achieved by inserting screws
radially into the outer cortex [31.80], however, if the
bone is osteoporotic, bone failure at the gripping site
may occur and confound the results. As an alternative, bones may be potted in an acrylic using such
products as Swiss Glas [31.81], Bondo [31.82], polymethylmethacrylate [31.83] or liquid metal [31.84].
Gripping Density
Bone strength is typically reported to be a function
of the square of the density, although there are reports that the strength can vary from a power of 1.3
to 3.0 [31.56]. Density varies geographically within
a body. Metaphyseal bone is more susceptible to resorption than cortical bone. There are also bones
which seem to be more affected by bone mineral
loss than others. In particular, the proximal femur
(hip), spine, and distal radius are common sites for
fragility (osteoporotic) fractures. Bone mineral density can be established nondestructively by means of
dual-energy x-ray absorptometry [31.86], quantitative
computed tomography [31.87], and ultrasound [31.88].
Alternatively, density can be obtained destructively by
removing (biopsy) a volume of bone, placing it in an
oven to eliminate all moisture, and then measuring the
resulting ash weight [31.89]. A less desirable method is
to obtain a standard x-ray in the field of view of which
is placed a step reference [31.90]. While all of the above
methods are valuable for documenting density and provide an indication of specimen strength, density is not
a good predictor of fracture strength. This is also true
clinically. A patient with low bone mineral density is at
risk of fracture, but density alone cannot predict with
any certainty when or if the fracture will occur.
Orientation
Bone is a transversely orthotropic material. When
measured along its longitudinal axis, bone exhibits an
ultimate strength of approximately 133 MPa in tension,
193 MPa in compression, and 68 MPa in shear [31.91].
879
Femoral bone is strongest along its longitudinal axis and
less strong in the transverse directions [31.91]. Similar
trends are noted for modulus. Actually, there is substantial variation in the reported strength and modulus
of bone [31.56]. The variation can often be attributed
to a multitude of factors that affect bone material behavior. These factors include the location, temperature,
orientation, hydration, gripping, and testing rate.
Viscoelasticity
Bone exhibits viscoelastic behavior; therefore, the rate
at which tests are conducted can dramatically affect the
measured response. Both the strength and stiffness increase as a function of increasing strain rate [31.92].
Storage
Bone is best stored wrapped in saline-soaked gauze
[31.93], double bagged, and frozen. Although it is preferred that tissue be tested as soon after harvesting as
possible, it is not always practical to do so. Often investigators need to thaw, prepare, refreeze, thaw, and
then test the specimens. Up to five cycles of freezing
and thawing of cortical bone reportedly do not alter the
compressive properties of bone [31.94]. For short-term
storage, −20 ◦ C seems sufficient. If tissue is to be stored
for the long term, −80 ◦ C freezer storage is preferred as
it minimizes enzymatic activity [31.93, 95].
31.5.2 Whole Bone
Whole-bone testing is used to measure the response
of a particular bone under typical in vivo loads. It is
often not possible, practical or ethical to acquire response parameters in vivo (particularly in humans), so
cadaveric tissue is commonly used as a model. Understanding the mechanism of injury requires knowing
the behavior of native tissue and how that behavior
changes as a function of age, loading/deformation rate,
healing, remodeling, drug use, and state of disease.
Another need for whole-bone testing is to investigate
the effect of stress concentrations on the structural response. In surgery, the placement of bone screws and
pins [31.96], as well as defects left after tumor resection [31.97, 98], can significantly weaken bone until
healing occurs [31.99]. Questions regarding the magnitude of the stress concentration and the duration of the
concentration (unlike engineering materials, bone usually restores itself to nearly initial values) are important
clinically [31.100].
Bones are normally not loaded in axial tension.
Even long bones (femur, humerus, tibia), which are
Part D 31.5
Environment
Bone is largely composed of water and its material
properties change as a function of hydration. For longterm tests in which the bone may be exposed to the
atmosphere, the bones must be hydrated [31.85]. Issues regarding hydration are similar to soft tissues. The
reader is also referred elsewhere to obtain a review of
general considerations of mechanical testing of bone
such as pH balance, temperature, tonicity, and the use
of antibiotics to prevent putrification during long tests.
31.5 Bone
880
Part D
Applications
Part D 31.5
nominally cylindrical but in reality have some curvature, are difficult to grip and fixture such that tension
is created uniformly across the cross section. For these
same reasons, it is difficult to load to failure bones in
tension without some bending superimposed by geometry or loading eccentricity.
Torsion tests are typically conducted on whole
bones to obtain shear properties [31.101, 102], but because of the natural curvature of long bones, it is often
difficult to apply only torsion.
Bones are naturally loaded in compression in vivo,
so it is not surprising that bone is strongest in compression. Loading in compression to obtain in situ strain
measurements is straightforward, but loading to failure
results in fractures along the shear planes of the bone.
31.5.3 Constructs
Construct testing is often used to compare various
modalities of fracture fixation and thereby provide advice for orthopaedic surgeons on the fixation that is
most stable or best satisfies biomechanical considerations (Fig. 31.5). Construct testing is often conducted
to gather basic performance data for use in the Food
and Drug Administration (FDA) approval process. Unfortunately, what constitutes an optimal construct is not
obvious or well documented. Interpreting the results of
a test and placing them in a clinically significant context
is perhaps the most challenging. In the 1960s the watch
words were rigidity of fixation. The prevailing wisdom
was to fix fractures as rigidly as possible and to place
screws in as many holes in the bone plates as were available. Surgeons quickly learned that, if the fixation was
excessively rigid, the plate stress shielded the fracture
site. As a result the osteoblasts did not receive the appropriate mechanical signals for full healing, that is, the
fracture site callus that formed did not ossify. Placing
bone screws in all the available holes was also unnecessary [31.103]. The screws closest to and furthest from
the fracture site are the ones that provide plate fixation.
Surgeons often place a third screw on each side of the
fracture as a safety in the event that one of the other
screws fails. The goal of fixation is to keep fracture site
motion to compressive strains below 2% [31.56]. Between 2% and 10%, a fibrocartilage callus forms and
places the fracture at risk of nonunion. There is a broad
range of construct stiffness that will yield the requisite
fracture site strains for healing [31.104].
Similar healing criteria are not available by which to
predict the efficacy of a particular construct based on the
mechanical response of the construct. For example, in
the spine, it is unknown how much motion spinal instrumentation should allow. If the instrumentation results
in too stiff a construct, the patient’s range of motion
may be impeded, degeneration of the disc proximal to
the top of the instrumentation, or fracture of the proximal vertebral level may occur. If the instrumentation
allows too much motion, the desired stabilization effect
of the instrumentation may not occur. The difficulty is
in determining the bounds of what constitutes optimal
fixation. The determination of such criteria is further
complicated by variation in patient activity level, bone
health (density), and changes in mechanical demands
placed on the instrumentation as healing occurs.
31.5.4 Testing Surrogates for Bone
Synthetic bones that have material properties similar to
native bone have been developed [31.105–108]. None
of these surrogate bones match all of the material
characteristics found in native bone [31.109, 110]. For
example, while some surrogate bones may match the
compressive modulus of normal bone, they may not
match the shear modulus. Some may exhibit appropriate quasistatic behavior, but when loaded at high strain
rates, the surrogate may not exhibit the appropriate viscoelasticity [31.111].
31.5.5 Outcome Measures
Fig. 31.5 Displacement measured on a femoral stem prosthesis relative to the proximal femur. An acupuncture
needle was inserted through the bone and points to a microscope calibration disc glued to the prosthesis
Kinematic
Linear variable displacement transducers (LVDT)s have
been mounted across the fracture site to measure fracture site motion, and if the initial fracture gap is known,
Experimental Methods in Biological Tissue Testing
A
PA
D
PB
P
B
B
Fig. 31.6 Ghost point P is a virtual point defined on A and
B. The virtual displacement of PB relative to PA is the true
displacement at the fracture site between fragment A and B
The analysis systems typically report rigid-body rotations and translations relative to the origin of a given
fragments reference system. If the origin coincides
with the rigid body’s center of rotation, translation
measurements can be decomposed into true translations and apparent translation due to rotation about
the instantaneous center of rotation. Unfortunately,
the center of rotation is not a fixed point for most
fragment motion [31.113]. One could define a fixed
reference, centered at a reproducible origin, i. e., center of mass, anatomical landmark, but such points are
hard to identify in biological tissues or are not practical. One solution is to establish ghost or virtual points
(Fig. 31.6). These are points of interest that are defined
relative to both fragments’ reference frames before testing begins. For example, if we are interested in fracture
gap motion, one could identify a point on the fracture
line and define that point in both fragment coordinate
systems. As the fragments move in rigid-body motion
relative to each other, the virtual location of the ghost
point can be determined in each fragment’s reference
frame and then mapped to the global reference frame.
The difference in the location of that ghost point, as predicted from the reference frame of each fragment, is the
true translation of one fragment relative to the other at
the location of the ghost point.
Perhaps the most popular of the motion analysis
systems use optical markers. These markers are either passive (reflective) markers or active light-emitting
diodes (LEDs). The accuracy of the LED systems is
reportedly [31.114] 0.3 mm in translation and 0.7◦ in
rotation. For passive markers, accuracy of 0.1 mm and
0.2◦ has been reported [31.81]. Active markers have
the advantage of being unambiguously recognized by
the receiver. Because the LED markers emit light in
a known sequence, the receiver can identify which
marker is active at any given point in time, whereas
passive markers have to be identified by their location
in the temporal context. The accuracy of the systems
is a function of the diagonal of the calibrated space,
so it is important to calibrate only the volume required
to conduct a given experiment, thereby maximizing the
system accuracy [31.115].
LEDs, because they are hard-wired to the recording system, can be cumbersome because of the multiple
wires running to the LEDs. Reflective markers are prone
to contamination with blood and lose their reflectiveness. Both types of markers can be obscured by the
testing apparatus and/or instrumentation. Further, optical systems using reflective markers are sensitive to
errant reflections from liquid (hydrated tissue) and shiny
881
Part D 31.5
calculate fracture site strain. Often the resulting fracture
site motion, even under a simple load (i. e., uniaxial) is
not one dimensional, making LVDT displacement difficult to interpret. Often the displacement sensed by
the LVDT is a combination of translational and rotational and impossible to decompose. The problem
has been overcome by mounting several LVDTs in
complex arrangements to calculate three-dimensional
(3-D) motion, but these methods are very cumbersome [31.112].
To overcome the shortcomings of LVDTs, motion
analysis techniques have been developed. These technique usually consist of rigidly mounting some type
of markers to the bone fragments and then placing the
fragments into a calibrated space. The location of the
markers is recorded by two or more receivers (cameras).
The relative location and angle of the receivers is known
(usually determined during calibration) so the 3-D coordinates of each marker location can be calculated by
triangulation. If the motion to be measured is known
to be planar, two-dimensional (2-D) coordinates can be
calculated using only one camera.
A major consideration of using the motion analysis systems is tracking motion at the site of interest.
31.5 Bone
882
Part D
Applications
Part D 31.5
surfaces such as polished stainless-steel orthopaedic instrumentation (bone plates, screw heads) or the columns
on the servohydraulic testing machine and fixturing.
Instead of using infrared light as the conduit for
obtaining position data, the flock of birds system
(Ascension Technology Corporation, Burlington, VT)
tracks the position and orientation of (slave) markers
based on the magnetic field strength of the master sensor [31.116].
Acoustic sensors function on a similar basis to optical systems except that markers consist of spark emitters
and the receivers are microphones [31.117]. Knowing
the speed of sound and the relative position of the microphones, the transit time of the sound pulse emitted
by a given spark is measured and the position of the
emitter can be triangulated. While accurate, these systems are prone to interference from ambient noise and
reflection off fixtures and test equipment.
Gross strains, such as the strain across a fracture
gap, can be calculated based on the displacements of
points across the fracture divided by the initial gap.
More typically we want measurements of strain in the
cortex of the bone. In this case, strain gauges can be
mounted directly to the cortex [31.118]. This requires
stripping of the periosteum, which serves as a vapor
barrier, so specimen dehydration must be considered.
In trabecular bone, digital image correlation has
been used to calculate strain [31.119].
Kinetic
Gross applied force may be measured at the point of
application using standard load cells. Because of the frequent use of baths and hydrated environment chambers,
load cells are often mounted to the actuator rather than
to the base of the testing apparatus, in order to prevent
damage to the electronics from saline. For high-speed
and cyclic tests, mounting the load cell to the actuator
requires that any mass distal to the sensing beams be
inertially compensated.
We often want to know what forces are transmitted
across fracture or osteotomy sites. Such measurements
provide information about the reduction provided by
various fixation devices. Changes in these measurements that occur with testing (acute or cyclic) inform
us as to how the fixation is behaving, i. e., if the instrumentation is loosening and how it is sharing load with
the bone.
Fig. 31.7 Washer load cell used to measure reduction
forces
A simple example of such a measurement is the use
of a washer load cell to measure the reduction force
provided by different types of lag screws for scaphoid
fixation Fig. 31.7. A more elaborate example is the use
of instrumented hardware and telemetry [31.120, 121].
While the data obtained from such devices is limited,
it gives us considerable insight into the loads transmitted during healing, partial weight-bearing, and activities
of daily living. The information is used to design rehabilitation programs, design new implants, and verify
computer models.
Some applications do not lend themselves to the
use of a washer load cell to measure contact force
as the washer may be too thick and thereby alter the
load transmission path or pressure distribution. In the
above example, pressure-sensitive film [31.122] could
have been employed, but the film only provides maximum pressure readings, so loss of reduction force
would have been missed. An alternative could be to use
a pressure-sensitive polymeric film [31.123]. While this
transducer provides real-time pressure distributions that
can be integrated to calculate the total contact force,
the transducers are temperamental, sensitive to temperature changes, need to be calibrated in the pressure range
in which they will be used, and are easily damaged if
kinked or exposed to sharp edges or to moisture, especially saline. The transducers are valuable for obtaining
relative pressures and spatial distributions of pressure.
They are less valuable for obtaining actual pressures
accurately.
Experimental Methods in Biological Tissue Testing
31.6 Skin Testing
883
31.6 Skin Testing
31.6.1 Background
31.6.2 In Vivo Testing
In vivo testing avoids the issues of tissue environment
and release of in vivo skin tension, and has the advantage of diagnosing skin dysfunction and pathology.
There are several difficulties of measuring mechanical
properties in vivo. Gripping the skin can be achieved
by gluing grips to the epidermis using cyanoacrylate
cement, but it is unknown how the surface traction is
transmitted through the depth of the skin. Skin thickness may be measured using ultrasound or skin fold
calipers, but both methods are prone to the inclusion of
some measurement artifacts.
Several attempts of uniaxial tests have been conducted by gluing grips to the skin surface. These tests
are affected by the thickness of the skin, but this can be
accounted for [31.130]. Of course uniaxial tests in vivo
are not truly uniaxial as the lateral boundaries are not
31.6.3 In Vitro
In vitro (ex vivo) testing reduces many of the challenges of in vivo testing, but replaces them with another
set of challenges. In vitro testing of excised specimens
permits the lateral boundaries to be traction free. The
tests may also be started in a stress-free state, however, it must be noted that the specimen is under some
in vivo tension that has been released during specimen
excision. The investigator must document the in vivo
specimen length so that, when it is excised and the specimen retracts, the investigator will know how to restore
the original dimensions. One of the simplest ways of
documenting in vivo dimension is to mark the skin with
a 2-D grid using an ink stamp [31.139].
Uniaxial specimens obeying dimensions of ASTM
D1708 can be stamped from skin specimens using a skin
punch [31.140]. The mechanical response of the uniaxial skin specimen is less than would be expected if the
specimen could be tested in vivo. Cutting uniaxial specimens from a field of skin severs the lateral boundary,
damages the collagen network, and isolates the specimen from the reinforcing effect of the neighboring
tissue.
Part D 31.6
Skin is the largest single organ in the body, accounting
for about 16% of the total body weight and has a surface
area of 1.5–2 m2 [31.2, 124]. Skin varies in thickness
from 0.2 mm on the eyelid to 6.0 mm on the sole of
the foot. Skin is composed of two layers: the epidermis
which forms the superficial layer that serves as a barrier, and the dermis which provides structural integrity
to the skin. As such, the dermis is the layer we are typically most interested in. The dermis consists mostly of
collagen fibers (type I and III), elastin, and reticulin, surrounded by a hydrated matrix of ground substance. Also
contained in the dermis are nerve endings, various ducts
and glands, blood vessels, lymph vessels, and hair shafts
and follicles.
The dermis has an upper layer (papillary) of fine,
randomly oriented collagen fibers [31.125] that connect the epidermis to the deep layer of the dermis,
the reticular dermis. The reticular dermis contains layers of thick, densely packed collagen fibers that are
organized in planes parallel to the surface of the
skin [31.126, 127] with some fibers traversing between
planes to limit interplanar shear. Within the planes, the
collagen fibers appear to have a preferential orientation [31.128, 129] that governs the anisotropic behavior
of skin. Surgeons are trained to cut along the dominant fiber orientation in order to minimize damage to
the fibers and reduce tension in the healing skin incision.
traction free, nor is the interface of the dermis with the
subdermis.
Some have tested the skin in torsion by attaching a disk to the skin’s surface and applying a known
twist or torque and measuring the response. The test
was modified by gluing an outer ring to the skin such
that only the annulus of skin between the ring and
the disk was tested [31.131]. Saunders [31.132] estimated the modulus of elasticity from torsional tests.
Wijn et al. [31.133] attempted to correlate torsional and
uniaxial test measurements using the theory of elasticity
of a homogeneous isotropic media, but did not recover
comparable material constants. They concluded that
skin cannot be treated as homogeneous and isotropic.
Troubled by the inability to retrieve material
constants, some investigators took a different approach [31.134–136]. They placed a type of suction
cup on the skin, applied a vacuum, and measured
the resulting dome height of the skin. A grid was
applied to the skin before the test to track 2-D deformation and it was found that the resulting fields were
inhomogeneous [31.137]. The investigators developed
a technique to start the tests with the skin in a stress-free
state [31.138].
884
Part D
Applications
Lanir and Fung [31.141] developed a device to test
membranous soft tissue biaxially. The device was capable of stretching in both directions, or one direction
while holding the other dimension constant or stress
free.
Because skin is not a homogeneous isotropic material, there is concern over whether the standard
dumbbell-shaped specimens satisfy the requirements
of being uniaxial with a uniform stress strain field
in the gage area. In composite materials, the lengthto-width ratio can be many times greater than that
needed for an isotropic material. In biological tissues, it is often not possible to obtain specimens
with such aspect ratios and in these cases pure
shear samples [31.142, 143] may be an attractive option.
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31.1
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31.2
31.3
31.4
31.5
31.6
31.7
31.8
31.9
31.10
31.11
31.12
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