Microcirculatory function monitoring at the bedside - A view from the intensive care

Abstract

Microcirculatory dysfunction plays a key role in the pathophysiology of various disease states and may consequently impact patient outcome. Until recently, the evaluation of the microcirculation using different measurement techniques has been mostly limited to animal and human research. With technical advances, microcirculatory monitoring nowadays becomes more and more available for application in clinical praxis. Unfortunately, measurements within the microcirculation are mostly limited to easily accessible surfaces, such as skin, muscle and tongue. Due to major differences in the physiologic regulation of microcirculatory blood flow and in metabolism between organs and even within different tissues in one organ, the clinical importance of regional microcirculatory measurements remains to be determined. In addition, technical methods available demonstrate large differences in the measured parameters and sampling volume, making interpretation of data even more difficult. Nonetheless, the monitoring of the microcirculation may, ahead of time, alert physicians that tissue oxygen supply becomes compromised and it may lead to a better understanding of basic pathophysiological aspects of disease. In the present review, we describe available non-invasive microcirculatory measurement techniques which can be applied clinically at the bedside. After a short discussion of physiologic and pathophysiologic basics related to microcirculatory monitoring, the measuring principles, applications, strengths and limitations of different monitoring systems are discussed.

Full text

IOP PUBLISHING PHYSIOLOGICAL MEASUREMENT
Physiol. Meas. 28 (2007) R65–R86 doi:10.1088/0967-3334/28/9/R01
TOPICAL REVIEW
Microcirculatory function monitoring at the
bedside—a view from the intensive care
Hans Knotzer
1
and Walter R Hasibeder
2
1
Department of Anesthesiology and Critical Care Medicine, Innsbruck Medical University,
Anichstrasse 35, A-6020 Innsbruck, Austria
2
Department of Anesthesiology and Critical Care Medicine, Krankenhaus der Barmherzigen
Schwestern Ried, Schlossberg 1, A-4910 Ried i.I., Austria
E-mail: johann.knotzer@uki.at
Received 17 March 2007, accepted for publication 25 June 2007
Published 21 August 2007
Online at stacks.iop.org/PM/28/R65
Abstract
Microcirculatory dysfunction plays a key role in the pathophysiology of
various disease states and may consequently impact patient outcome. Until
recently, the evaluation of the microcirculation using different measurement
techniques has been mostly limited to animal and human research. With
technical advances, microcirculatory monitoring nowadays becomes more and
more available for application in clinical praxis. Unfortunately, measurements
within the microcirculation are mostly limited to easily accessible surfaces,
such as skin, muscle and tongue. Due to major differences in the physiologic
regulation of microcirculatory blood flow and in metabolism between organs
and even within different tissues in one organ, the clinical importance of
regional microcirculatory measurements remains to be determined. In addition,
technical methods available demonstrate large differences in the measured
parameters and sampling volume, making interpretation of data even more
difficult. Nonetheless, the monitoring of the microcirculation may, ahead of
time, alert physicians that tissue oxygen supply becomes compromised and
it may lead to a better understanding of basic pathophysiological aspects
of disease. In the present review, we describe available non-invasive
microcirculatory measurement techniques which can be applied clinically at the
bedside. After a short discussion of physiologic and pathophysiologic basics
related to microcirculatory monitoring, the measuring principles, applications,
strengths and limitations of different monitoring systems are discussed.
Keywords: gastric tonometry, laser Doppler flowmetry, venous
occlusion plethysmography, near-infrared spectrophotometry, tissue reflectance
spectrophotometry, orthogonal polarization spectral imaging
(Some figures in this article are in colour only in the electronic version)
0967-3334/07/090065+22$30.00 © 2007 IOP Publishing Ltd Printed in the UK R65

R66 Topical Review
1. Introduction
Microcirculatory dysfunction has been hypothesized to play a key role in the pathophysiology
of organ failure and consequently patient outcome (Lehr et al 2000, Dhainaut et al 2005).
Therefore, different measurement technologies focusing on the microcirculation were
applied to critically ill patients in order to detect microcirculatory failure at an early stage.
Clinical investigations, including limited numbers of patients, have provided some evidence
demonstrating that microcirculatory dysfunction can be recognized early and that specific
treatment may improve tissue oxygen supply (De Backer et al 2006, Morelli et al 2004).
Unfortunately, available technology still demonstrates substantial drawbacks. Application
of sensors is mostly limited to superficial tissues and measurements are performed within small
tissue volumes at one site. Thus, regional measurements in one organ may not be representative
for other organs or even for different tissue compartments in the same organ.
Furthermore, techniques which pretend to measure microcirculatory function are quite
heterogeneous in the parameters they are assessing (table 1). For example, blood flow
measurements can be performed over more or less clearly defined tissue samples resulting in
a signal originating from arterioles, capillaries and venules as in the case of laser Doppler
flowmetry. On the other hand, microcirculatory blood flow can be directly visualized and
disturbances related to distinct vascular segments using intravital microscopic devices. Other
methods, e.g. near-infrared spectroscopy, polarographic oxygen sensors, microcirculatory
stress tests, e.g. the reactive hyperemia response, and venous congestion plethysmography
only indirectly assess parameters related to microcirculatory function. Furthermore, one
could argue that some of these techniques, such as near-infrared spectroscopy or polarographic
oxygen sensors, although not directly measuring blood flow, more accurately reflect the effect
of microcirculatory blood flow disturbances due to tissue oxygen supply alterations compared
with quantitative measurements of blood flow. These facts and considerations are making the
interpretation of results even more difficult.
In the present review, we will first discuss some physiologic and pathophysiologic basics
concerning the regulation of organ blood flow, which have major impact on microcirculatory
monitoring and correct data interpretation. Second, we introduce various measurement
techniques currently applied for the assessment of the microcirculation in critically ill patients.
The principles of measurements and the techniques will be briefly discussed with respect to
their possible advantages and limitations.
2. Physiologic and pathophysiologic basics as related to microcirculatory monitoring
For a long time it has been known that global measurements of systemic hemodynamic
parameters, systemic oxygen transport variables or serum lactate concentration may not
necessarily reflect the adequacy of tissue oxygen supply at the microcirculatory level. However,
from the standpoint of physiology, ‘normal’ hemodynamic conditions are a precondition for
regular microcirculatory function in tissues. Vascular conductivity, e.g. the reciprocal of
vascular resistance, varies widely between organs, and is regulated by metabolic, hormonal
and neuronal factors (Rowell 1986, Shipley and Study 1951, Lassen 1959, Feigl 1983, Granger
et al 1980). The influence of hormones and the sympathetic nervous system on organ blood
flow heavily depends on receptor types, receptor density and distribution especially in arteriolar
vessels (Guimaraes and Moura 2001). Table 2 summarizes basic physiologic data derived from
animal experiments on receptor types, receptor density, the relative strength of blood flow auto
regulation, and the lower auto regulatory perfusion pressure, e.g. the blood pressure below
which organ blood flow becomes dependent on perfusion pressure.

Topical Review R67
Table 1. Classification of microcirculatory monitoring systems according to the main parameters measured.
Monitoring system Measuring principle Measured parameters Application Advantage Disadvantage
Orthogonal polarization Visualization of the Vascular diameter, Sublingual, Short measuring period, Susceptible to pressure
spectral imaging (OPS) microcirculation with blood flow velocity, skin in newborns, assessment of regional and movement artifacts,
polarized light with a functional capillary surface of internal organs heterogeneity in blood time expensive and
wavelength of 548 nm density, in the operating theatre flow semiquantitative off-line
determination of low analysis,
flow and intermittent small penetration depth
flow of 1 mm
Venous occlusion Measurement of Blood flow, Skeletal muscle of the Absolute values Time expensive and
plethysmography circumference changes of vascular permeability, extremities complex calibration,
the limb after increasing balance of the Starling susceptible to motion
the hydrostatic pressure forces artifacts,
false interpretation of
filtration coefficient in
patients with low
diastolic blood pressure
pH
i
-tonometry Equilibration of an air- or Intraluminal pCO
2
Sublingual, Easy to use No consideration of
fluid-filled balloon tipped arterial–mucosal pCO
2
- stomach, regional perfusion
catheter in the stomach gap, gut heterogeneity,
with pCO
2
originated pH
i
interference with gastric
from mucosal tissue acid, enteral feeding,
duodenal reflux
Laser Doppler Quantification of Flux of erythrocytes All organ surfaces, Easy handling, Small penetration depth
flowmetry backscattered Doppler- especially skin and short measuring period of about 1 mm,
shifted light from tissue gastric mucosa no absolute values,
in motion motion detection without
blood flow = biological zero

R68 Topical Review
Table 1. (Continued.)
Monitoring system Measuring principle Measured parameters Application Advantage Disadvantage
Tissue reflectance Detection of Hemoglobin oxygen All organ surfaces, Easy handling, Wavelength spectrum is
photometry backscattered light with saturation, especially skin and gut absolute values, influenced by other tissue
specific wavelength capillary hemoglobin mucosa short measuring period chromophores (e.g.
spectra (oxygenated concentration melanin and
hemoglobin two peaks at cytochromes) and path length
542 and 577 nm; of the photon
deoxygenated through different tissue,
hemoglobin one peak 556 nm) detection of mainly
capillary-venous oxygen
saturation (85%)
Near-infrared Infrared light (700– Tissue hemoglobin All organ surfaces, Easy to use, Inability to measure
spectroscopy 1000 nm) penetrates the oxygenation, especially brain through short signal response absolute tHbO
2
values,
tissue, is absorbed by redox status of the skull and muscle time unknown exact
chromophores cytochrome aa3 tissue penetration depth of light,
(hemoglobin, myoglobin poor spatial resolution of
and cytochrome aa3), the signal
and mean tissue
hemoglobin oxygen
saturation (tHbO
2
)is
calculated based on the
Beer law

Topical Review R69
Table 2. Physiologic data of receptor types, receptor density, the relative strength of blood flow autoregulation and the lower autoregulatory perfusion pressure in various organs.
Organ Receptors Autoregulation Lower autoregulation threshold Remarks
Brain α
1A
± ++++ ∼50 mmHg Only transient minor vasoconstriction to SNS
α
1D
±
Heart α
1A
+(+) In normal hearts only very high SNS may cause vasoconstriction
β
1
+++ High AVP may cause vasoconstriction
(β
2
) (+) ++++ ∼40 mmHg Moderate AVP vasodilation
V
1A
+ Excessive catecholamine stimulation may cause ischemia and
regional inflammation
Kidney α
1A
(+) SNS can transiently decrease RBF as low as 20–30% of normal
α
1B
++ Infusion of NE <µgkg
−1
min
−1
i.v. does not induce ATN
(β) (+) ++++ ∼60–70 mmHg V
1A
mainly located at glomerular efferent vessels
V
1A
+
V
2
++
GIT α
1A
+++ Autoregulation depends on feeding status
α
1B
+
α
1D
+ ∼40 mmHg
β
2
+++ ++
V
1A
++
Muscle α
1A
(+) Autoregulation prevails only during exercise
α
1B
+++ + At rest ??
β
2
+++
V
1A
++
Skin α
1A
+++ Virtually no autoregulation of blood flow
α
2
+– –
V
1A
+++
Peripheral nerves α ± – – Virtually no autoregulation of blood flow
V, arginine–vasopressin receptors; SNS, sympathetic nervous system; AVP, arginine–vasopressin; RBF, renal blood flow; NE, norepinephrine; ATN, acute tubular necrosis.
(+) (probably) weak action.
++ moderate action.
+++ strong action.
++++ very strong action.

R70 Topical Review
Differences in receptor types, receptor density and the importance of blood flow
auto regulation, which principally indicates the strength of the metabolic component of
microcirculatory regulation, explain many of the problems regarding the correct clinical
interpretation of data obtained from microcirculatory monitoring. For example, any stress
response or infusion of vasoactive drugs may acutely decrease microvascular blood flow in the
skin and muscle because of the high density of arteriolar α-adrenoreceptors and V
1
-arginine
vasopressin receptors and the almost absent microcirculatory metabolic blood flow control at
rest. However, ‘internal organs’, e.g. gastrointestinal tract, kidney or heart, possess strong
vasodilating mechanisms and exert a much more pronounced metabolic control of blood flow
when compared with skin or skeletal muscle. Thus, microcirculatory blood flow in these
organs may be well maintained. In addition, since for example the skin and muscle have very
low basal metabolic rates for oxygen, any moderate disturbance of regional oxygen supply will
unlikely be of clinical importance. Other, more easily accessible sites for microcirculatory
measurements, such as the eye or the sublingual mucosal area, are possibly better clinical
indicators for microcirculatory responses under pathophysiological states and vasoactive
therapy. For example, the retina, similar with the brain, is an organ with high metabolic activity
(Haefliger et al 2001, Faraci 1992). Autoregulation plays an important role in these structures
of high metabolism, and the balance between vasoconstricting mediators, e.g. endothelin, and
vasodilators, e.g. nitric oxide, is very vulnerable, leading to disturbances in blood supply
to the tissue (Haefliger et al 1994). The sublingual mucosal area gets more attention as a
monitoring site for microcirculatory alterations, as the sublingual mucosal microcirculation
shares the embryologic origin with the digestive mucosa, a mucosa highly vulnerable
to systemic hemodynamic and microcirculatory deterioration. A recent investigation of
Creteur et al could demonstrate a significant correlation between sublingual mucosal carbon
dioxide pressure, microvascular flow index and gastric mucosal carbon dioxide pressure in
septic patients (Creteur et al 2006), indicating an interesting window for microcirculatory
monitoring.
Under pathophysiologic conditions, e.g. septic shock, general hypoxia, ischemia-
reperfusion injury, intense activation of the immune system, vascular endothelial cells, the
clotting system and direct tissue injury may have pronounced effects on the microcirculation.
Endothelial swelling, leucocyte plugging, disseminated intravascular clot formation, tissue
edema and the build up of vasoactive mediators (e.g. nitric oxide, arachidonic acid metabolites,
histamine, serotonine, adenosin) may lead to increasing heterogeneity of microvascular blood
flow, thereby decreasing tissue oxygen supply and promoting organ failure (Ellis et al 2002,
Mazzoni et al 1989, Talbott et al 1994, Eichelbronner et al 2003, Gando et al 1999,
Habazettl et al 1999) (figure 1). In addition, systemic hemodynamic failure due to down-
regulation of adrenergic receptors on vascular smooth muscle cells, depletion of endogenous
vasoactive hormones (e.g. cortisol, arginine vasopressin), and increased synthesis and release
of vasoactive substances (e.g. arachidonic acid metabolites, nitric oxide, adenosine), may
significantly contribute to perfusion heterogeneity (Spain et al 1999, Annane et al 1998,
Dunser et al 2003,Tymlet al 1998, Messina et al 1988, Nishiyama et al 2004).
Unfortunately, the time course and the exact contribution of microcirculatory failure
to the development of multiple organ dysfunction syndrome in patients still remain to be
determined. Recent studies suggest that cytopathic hypoxia, due to mitochondrial dysfunction
and increased apoptosis, may also have a significant impact on patient prognosis (Brealey
et al 2002, Papathanassoglou et al 2000). Furthermore, beyond pure manipulation of systemic
hemodynamics, specific therapy to protect or improve the microcirculation is still experimental,
and therefore data are mostly limited to short time animal models of defined pathophysiology
(Schaser et al 2005, Nolte et al 2004, Matejovic et al 2007).

Topical Review R71
Figure 1. Microvascular plugging with leucocytes and macrophages in the lungs of a patient
who died in severe multiple organ dysfunction syndrome due to acute necrotising pancreatitis with
formation of a large retroperitoneal abscess.
Despite these obvious restrictions in connexion with interpretation and therapeutic options
related to data obtained from microcirculatory monitoring techniques, measurements of
microcirculatory parameters may, ahead of time, alert clinicians of deteriorating tissue oxygen
supply. Therefore, clinicians caring for severely ill patients should stay engaged with new
technical advances in microcirculatory monitoring and closely watch new therapeutic options
to improve microcirculatory function.
3. Microcirculatory measurement techniques
3.1. pH
i
-tonometry
The gastric or intestinal luminal pH-tonometry (pH
i
-tonometry) is used as an indirect method
to assess the adequacy of oxygen supply to the gastrointestinal mucosa (Cerny and Cvachovec
2000). The measurement technique is based on the determination of pCO
2
in the gut luminal
fluid with a balloon tipped catheter (figure 2). For the calculation of pH
i
, it has to be assumed
that mucosal intracellular pCO
2
is in equilibrium with the luminal fluid and the semipermeable
saline or air-filled tonometer balloon and that blood bicarbonate equals mucosal bicarbonate
concentration. With these assumptions intramucosal pH can be easily calculated with the

R72 Topical Review
Figure 2. Schematic diagram of gastric tonometry. Hydrogen ions are buffered by tissue
bicarbonates resulting in CO
2
production.
Henderson–Hasselbach equation (Fiddian-Green et al 1982)
pH
i
= 6.1+log
10
[HCO
−
3
]
α p
r
CO
2
= 7.37 ± 0.04 (Gutierrez and Brown 1996),
where pH
i
is the calculated intramucosal pH, 6.1 is the pK
a
of carbonic acid, [HCO
−
3
]is
the concentration bicarbonate in mmol l
−1
of the mucosa which is assumed to be equal to
arterial concentration of bicarbonate, α is the Bunsen solubility coefficient, which represents
the solubility of CO
2
in plasma (0.0301) at 37
◦
C, p
r
CO
2
is the measured intraluminal carbon
dioxide tension in kPa, which is assumed to be in equilibrium with the intramucosal pCO
2
.
Unfortunately, the assumption made for the calculation of pH
i
may not be valid during
various shock states and systemic acid base disorders (Vincent and Creteur 1998). During
gastrointestinal low flow conditions mucosal bicarbonate may decrease more rapidly than
arterial bicarbonate (Antonsson et al 1990). In systemic acidosis, low pH
i
values may purely
reflect generalized acidosis and not local tissue hypoxia. Considering these difficulties in
the interpretation of pH
i
, the use of the gastric–arterial pCO
2
difference (pCO
2
-gap) has been
proposed as the more appropriate tonometrically derived variable for interpretation (Vincent
and Creteur 1998).
Two mechanisms are involved in the generation of regional tissue pCO
2
. On the one
hand, increased tissue pCO
2
may reflect low mucosal blood flow at constant metabolism with
diminished evacuation and consequently accumulation of CO
2
in the gastrointestinal lumen.
On the other hand, anaerobic metabolism with a net hydrogen ion production may increase
tissue pCO
2
H
+
+HCO
−
3
↔ H
2
CO
3
↔ H
2
O+CO
2
.
Therefore, systemic acid–base status must always be evaluated before the final interpretation
of tonometrically derived variables.
Various methodological problems exist in the clinical use of pH
i
-tonometry. First, the
use of saline-filled balloon catheters for tonometry may underestimate regional pCO
2
(Takala
et al 1994). This problem could be solved by using phosphate buffered solutions (Knichwitz
et al 1996). Because of equilibration periods exceeding 90 min, this method has limitations

Topical Review R73
in the clinical setting. Second, enteral feeding may stimulate the secretion of hydrogen ions
by parietal cells of the mucosa. The H
+
ions may interact with HCO
3
−
in the mucosal layer
and increase tonometer pCO
2
. For that reason, a discontinuation of enteral feeding for at
least 60 min before measuring and the application of H
2
blockers are recommended to reduce
possible bias in tonometry measurements (Marik and Lorenzana 1996, Kolkman et al 1994).
Third, reflux of HCO
−
3
-rich duodenal fluid in the stomach can also increase regional pCO
2
(Fiddian-Green et al 1982).
Despite several physiological assumptions, methodological problems and the ongoing
discussion concerning the origin of increased mucosal pCO
2
, low gastric mucosal pH
i
was
reported as a risk factor associated with increased mortality in critically ill patients during
the first 12 h after admission (Doglio et al 1991). Similar results were confirmed in the later
publications, focusing on the detection of splanchnic hypoperfusion and on the prediction of
mortality (Maynard et al 1993,Marik1993). In addition, patients with a pH
i
less than 7.32
after major surgery had a longer stay on ICU, increased incidence of major complications,
higher mortality rate and higher costs (Mythen and Webb 1994). Similarly, an increased
pCO
2
-gap, representing the overall balance of oxygen supply to demand within the gastric
mucosa, has been associated with an unfavorable outcome in ventilated patients (Levy et al
2003).
Other investigators failed to demonstrate a link between tonometrically derived variables
and outcome in critically ill patients (Gomersall et al 1997). In hemodynamically stable,
resuscitated surgical intensive care patients, we found that pH
i
and pCO
2
-gap did not reflect
severity of multiple organ dysfunction syndrome (Knotzer et al 2006). In another study, the
resuscitation of critically ill patients based on gastric tonometry failed to improve patients
outcome (Gomersall et al 2000).
3.2. Laser Doppler flowmetry
Laser Doppler flowmetry (LDF) is a non-invasive instrument permitting real-time measurement
of microvascular perfusion, particularly in the skin. Monochromatic laser light with a
particular wavelength penetrates the surface of a sampling volume and interacts with both
static and moving cells (i.e. red blood cells). Due to the Doppler effect, photons of the
laser light scattered on blood cells undergo a frequency shift proportional to the speed of
the moving cells. Backscattered light is transmitted via a flexible optical fiber to the laser
Doppler photodiode, amplified, analyzed, and finally transformed into an analog signal. The
magnitude and frequency distribution of changes in wavelength are proportional to the number
of blood cells multiplied by the mean velocity of these cells. This product is termed flux and
is proportional to flow.
A wide range of LDF probes is available for different measurement sites. The measuring
depth depends on the wavelength used and on the distance between transmitting and receiving
fibers of the probe (Obeid et al 1988). Most Doppler devices use a fiber separation of 500 µm
resulting in a penetration volume of about 1–1.5 mm
3
.
In contrast to conventional LDF probes, laser Doppler imaging uses a moving mirror
to direct a 632.8 nm red laser beam directed to the skin. Backscattered light from each
measurement point is detected separately, and a two-dimensional color-coded image of skin
perfusion is generated. An advantage of laser Doppler imaging is the observation of a
larger tissue area and a better spatial resolution of tissue blood flow (Forrester et al 1997).
A disadvantage is the data evaluation of the two-dimensional pictures, calculating a mean
perfusion from the ‘perfusion map’.
Unfortunately, LDF does not record absolute blood flow values, e.g. in ml × min
−1
×
mm
3
. With LDF tissue perfusion is expressed in terms of arbitrary perfusion units (PU),

R74 Topical Review
where 1 PU represents a pre-defined electrical signal in mV. Under experimental and clinical
conditions a close relationship between LDF derived and directly measured blood flow exists
(Johnson et al 1984). Recently, a more sophisticated approach to express LDF data has
been advanced dividing LD-derived blood flow measurements by mean arterial blood pressure
(Cracowski et al 2006). This calculation gives the ‘relative’ vascular conductance, a measure
of the blood flow in direct dependence to the actual perfusion pressure in the observed tissue.
Substantial spatial and temporal variations in LDF-derived blood flow measurements may
exist, in particular during repeated examinations (Tenland et al 1983). LDF recordings are
affected by regional microcirculatory blood flow heterogeneity, differences in environmental
conditions, e.g. skin temperature, and changing of the recording location (Stucker et al 2001).
If the recording location is standardized, the mean coefficients of variation of LDF-derived
blood flow measured on different days in the same individual are <10% (Kubli et al 2000).
A novel laser Doppler related technique is the laser speckle contrast analysis (LASCA),
which provides a high-resolution blood flow imaging method (Briers 2001). Laser speckle is a
random interference pattern which occurs when coherent laser light is scattered from a tissue.
Due to minor differences of the traveled path length, the light rays differ in phase and can
thus interfere in a constructive or destructive manner. Each image point is subject to intensity
fluctuations that depend on these phase differences. The image of a laser-illuminated surface
thus appears granulated, the so-called speckled. Moving particles within the tissue, e.g. red
blood cells, change the scattering properties dynamically and produce a time-varying speckle
pattern in each pixel of the image. If the blood flow increases, the intensity variation of the
speckle pattern is more rapid, which leads to a loss of contrast. With a LASCA device, these
varying speckle patterns are quantified and analyzed (Dunn et al 2001). Light from a laser
is diverged by a lens to illuminate the area of tissue under investigation. A camera images
the illuminated area and the image is observed on a monitor. A specially developed software
processes it to produce a false-color contrast map indicating velocity variations.
LDF has been extensively used for blood flow measurements, in particular in the skin
and the gastrointestinal tract, under experimental and in clinical conditions (Haisjackl et al
1990, Knotzer et al 2006, Young and Cameron 1995, Luckner et al 2006). The penetration
depth of LDF in the human skin is about 1.0–1.5 mm. In the skin, nutritional capillaries are
most superficial (0.1–0.5 mm distance from skin surface). Therefore, only a small fraction
(5–10%) of skin blood flow passes through this superficial nutritional layer (Bongard and
Fagrell 1990). Deeper subpapillary tissue represents the thermoregulatory microvascular bed
(0.05–2.0 mm) with predominantly venular vessels (95%), and only few arterioles. Therefore,
LDF measurements in the skin represent predominantly thermoregulatory blood flow with
only a small fraction of ‘true’ nutritional flow.
The intestinal tract is assumed to play an important role in the development and
maintenance of multiple organ dysfunction in critically ill patients. Since the mucosa layer
represents the physiologic barrier against intestinal bacteria and their toxins, mucosal blood
flow has been a major focus of scientific interest. LDF is used to measure blood flow in
the gastric wall in critical ill patients (Duranteau et al 1999,Neviereet al 1996). Although
authors pretend that their measurements are limited to the mucosal layer, experimental data
suggest that LDF assesses blood flow of the whole gastrointestinal wall (Schwarz et al 2001).
Measurements of total backscattered light on an excised piece of pig jejunum demonstrated a
profound increase when a mirror was intermittently placed under the preparation proving the
fact that the laser beam penetrates the whole intestinal wall and is not limited to the mucosal
layer (figure 3).
With LDF the microcirculation can be functionally assessed using different stimulation
tests, e.g. the measurement of postocclusive reactive hyperemia, blood flow changes during

Topical Review R75
Figure 3. Amount of total backscattered (TB) light and perfusion units (PU) from a laser Doppler
flowmeter without and with a mirror intermittently placed under an excised part of jejunum. TB
and PU increased with a mirror under the preparation, showing that the signal of laser Doppler is
sampled throughout the whole intestinal wall.
local skin warming or as skin microcirculature response to iontophoresis (Knotzer et al 2006,
Cracowski et al 2006, Christen et al 2004). The major advantage of stress testing may be that
they provide a more meaningful way of assessing the microcirculation than just measuring
resting blood flow.
3.3. Venous occlusion plethysmography
Venous occlusion plethysmography (VOP) is primarily used to measure blood flow in humans.
The underlying principle is relatively simple: venous drainage from an extremity is briefly
interrupted with a pressure cuff while arterial inflow is initially unaltered resulting in a linear
increase in tissue volume, which is proportional to arterial blood inflow, until venous pressure
rises towards the occluding pressure (Greenfield et al 1963). In clinical practice volume
changes are detected with electromechanical strain gauge sensors automatically stretched to
a predetermined tension. Changes in resistance of the gauge result from alterations of limb
girth by increasing venous congestion upstream to the strain gauge (Gamble et al 1993).
If hydrostatic pressure raises further, a second phase characterized by a slow gain in limb
volume can be recorded (figure 4). This phase is due to increased filtration of fluid from
the microvasculature into the interstitium. During a continuous series of increasing defined
pressure steps the vascular compliance and the fluid filtration component of the increase in

R76 Topical Review
(a)
(b)
Figure 4. (a) Original recording of changes in limb volume resulting from a stepwise increase
(10 mmHg for 270 s) in cuff pressure for the determination of capillary permeability in a patient
with multiple organ dysfunction syndrome. (b) Volume response to a single cuff pressure increase
lasting 270 s for analysis. There is a rapid increase in limb volume according to venous distension,
followed by a slower continued rise in volume due to movement of fluid from the capillaries into
the limb.
limb circumference can be analyzed (Wilkinson and Webb 2001, Gamble et al 1993). At each
pressure step, the cuff pressure (mmHg) and the slope of the slow volume change indicating
increased capillary filtration (Jv; ml × 100 ml
−1
× min
−1
) are recorded. The values of
Jv, when plotted against corresponding cuff pressures show a linear correlation (Bauer et al
2002). The interception with the x-axis reflects the intravascular pressure (Pv), where capillary
filtration starts. The slope of the plotted line corresponds to the capillary filtration coefficient
(figure 4).
VOP is a highly interesting technique for investigations of the microcirculation because it
indirectly assesses vascular endothelial barrier function. Presumably, changes in microvascular
permeability are one of the foremost alterations occurring within the microcirculation during
pathophysiologic conditions, e.g. sepsis, ischemia-reperfusion injury (Langheinrich and
Ritman 2006, Pickkers et al 2005).
One major limitation of VOP in clinical practice is that measurements are limited to
extremities which mainly represent skeletal muscle tissue. Because of the existence of
major differences in capillary histology (e.g. fenestrated versus continuous or discontinuous
capillaries) microvascular permeability may not change uniformly during disease. However,
a recent study evaluating contrast agent diffusion across the vascular endothelium in an
animal model of lipopolysaccharide-induced sepsis demonstrated simultaneous changes in

Topical Review R77
microvascular permeability in the heart, kidney and colonic wall (Langheinrich and Ritman
2006).
Another limitation for a more widespread use of the VOP technique is the relatively
complicated clinical application. VOP requires long measurement periods exceeding 30 min
per measurement. During calibration and final measurement the patient has to rest motionless
in bed to avoid measuring artifacts. In critically ill patients this usually requires deep
analgosedation and sometimes relaxation. Finally, in patients with a very low diastolic
blood pressure and/or a high isovolumetric venous pressure, e.g. the pressure where fluid
filtration starts, the measurement protocol may not allow an accurate estimation of the filtration
coefficient, because fluid filtration is only observed only at high cuff pressures close to diastolic
blood pressure.
In clinical practice VOP has been successfully applied to determine microvascular
permeability in diabetic patients (Jaap et al 1993), patients with septic shock (Christ et al
1998), in response to infused peptides (Ando et al 1992), and in hemodynamically stable
patients suffering from multiple organ dysfunction syndrome (Knotzer et al 2006).
3.4. Tissue reflectance spectrophotometry
The first rapid tissue reflectance spectrophotometer (TRS) was introduced by Luebbers in
1957, and the hemoglobin spectra were presented in 1964 by the same study group (Luebbers
and Niesel 1957,Nieselet al 1964). TRS was further improved by Sato et al and by Frank
et al as a non-invasive method for assessing tissue hemoglobin oxygen saturation (tHbO
2
) and
changes in tissue hemoglobin concentration (tHbC) (Sato et al 1979, Frank et al 1989). The
light of a xenon high-pressure arc lamp is transferred to a tissue surface via a single highly
flexible micro-lightguide with a diameter of 250 µm. Backscattered light is collected by six
identical micro-lightguides (each 250 µm in diameter) arranged around the circumference of
the illuminating lightguide. From there it travels to a rotating bandpass interference filter disc,
which serves as a monochromating unit for a spectral range of 502–628 nm. This filter disc
permits sampling in steps of 2 nm and a sampling rate of 100 spectra s
−1
. The monochromated
light is transmitted to a photomultiplier tube, fed into a current-to-voltage converter and
amplified by a cascade amplifier. The voltage signal is offset compensated, filtered by a
low-pass filter, fed into an analog-to-digital converter and transferred to a computer. The
recorded spectra are balanced against a standard white reference produced by a mirror and a
dark reference. Absolute values of tHbO
2
and relative values of hemoglobin concentration
tHbC are calculated by using an algorithm, originally developed by D
¨
ummler and described in
detail by Frank et al (1989). This algorithm has been validated for mucosal tissue (Hasibeder
et al 1994). In brief, the algorithm is a derivation of the differential equations used in the
Kubelka and Munk theory (Kubelka and Munk 1931). This equation mathematically describes
the relation between the wavelength-depending absorption and the wavelength-depending
scattering of the tissue. This relation in turn depends on four parameters, the basic absorption
of the tissue, the basic scattering of the tissue, the concentration of oxygenated hemoglobin
and the concentration of deoxygenated hemoglobin. The hemoglobin oxygenation can be
evaluated due to re-formatting the equation and measuring the basic scattering of the tissue.
The back scattering of the light requires constancy for exact evaluation of the hemoglobin
oxygenation.
TRS has been extensively used in animal experiments in particular to assess the adequacy
of oxygen supply to the gastrointestinal tract (figure 5) (Knotzer et al 2005, Germann et al
1997, Hasibeder et al 1994). When used on the mucosal surface in pigs the light signal of
the TRS similar to laser light penetrates the whole gastrointestinal wall (Schwarz et al 2001).

R78 Topical Review
Figure 5. Original tracing of tHbO
2
(black line) and tHbC (gray points) obtained from the jejunal
mucosa of a pig with a tissue reflectance spectrophotometer (EMPHO II). The fast signal sampling
rate allows for resolution of regular changes in both signals originating from arteriolar vasomotion,
e.g. regular variations in arteriolar vasomotor tone. The regular wide changes in tHbO
2
and tHbC
are most likely the result of the countercurrent arrangement of microvessels in conjunction with
vasomotion.
Therefore, the measurement signal is not only limited to the mucosal layer although tHbO
2
values detected from the mucosa are lower than measured from the serosa suggesting that most
signal originates from the mucosal layer just beneath the sensor. Changes in relative tHbC,
which can be measured by TRS, track changes in systemic hematocrit (Haisjackl et al 1997).
However, during progressive anemia induced by isovolemic hemodilution, tHbC decreased less
in magnitude when compared with systemic hematocrit supporting the hypothesis that tissue
hematocrit is a more controlled physiologic variable during progressive anemia (Haisjackl
et al 1997).
Amongst other applications, TRS in humans has been widely used to assess gastric
mucosal tHbO
2
in septic shock, during and after coronary artery bypass graft surgery, and
during progressive PEEP application (Temmesfeld-Wollbruck et al 1998, Fournell et al 2002,
2003). The method is highly suitable to detect fast changes in tissue oxygen supply in the
gastric mucosa (Friedland et al 2004). However, the application is intermittent and the micro-
lightguides have to be applied via the channel of a gastroscope onto the mucosal surface.
Therefore, investigators need to have practice to avoid artifacts due to oversized pressure to
tissue and ongoing gastric contractions during measurements.
3.5. Near-infrared spectroscopy
Near-infrared spectroscopy (NIRS) represents a special form of tissue reflectance spectroscopy
(Jobsis 1977, Brazy et al 1985). Infrared light closest to the visual spectrum (700–
1000 nm) easily penetrates soft tissue and even bone and is absorbed by chromophores
including hemoglobin, myoglobin and cytochrome aa3. Light absorption within tissue
depends on the concentration and oxygenation status of chromophores, in particular oxy- and
deoxyhemoglobin. Like pulse oxymetry, NIRS calculates mean tissue hemoglobin oxygen
saturation (tHbO
2
) based on the Beer law relating light absorption (measured by the device) to
oxyhemoglobin and deoxyhemoglobin concentration in the tissue. However, exact calculations
require constancy of the optical path length and constant light scattering in the tissue, making
absolute measurements of tHbO
2
in clinical practice impossible. However, NIRS is suitable
to measure changes in tHbO
2
and thus tissue oxygen supply over time.
Currently, continuous wave near-infrared spectroscopic devices are the most widely used
instruments to monitor tHbO
2
in particular during carotid endarteriectomy, in patients with
brain injury, during cardiac surgery, and in preterm and term neonates (Mille et al 2004,
Ogasawara et al 2003, Brawanski et al 2002, McQuillen et al 2007, Wolfberg and du Plessis

Topical Review R79
2006). Concerning the brain, simultaneous measurements of arterial, jugular bulb and tHbO
2
suggest that the NIRS signal is derived from arterial and venous tissue compartments at a ratio
of approximately 16:84. This ratio was similar during normoxia, hypoxia and hypocapnia
(Watzman et al 2000). More recently, one study in late-term pregnant women reported the
possibility of measuring fetal brain oxygenation which ranged between 50% and 74% through
the maternal abdomen (Vintzileos et al 2005). However, currently there is no way to validate
these data.
Currently, NIRS seems to be a suitable instrument to measure changes in brain tissue
oxygen supply, by measuring tHbO
2
. NIRS devices available for clinical application are easy
to operate and have an adequate response time. The major disadvantage of continuous wave
NIRS is the inability to measure absolute tHbO
2
values, the unknown exact penetration depth
of light and the poor spatial resolution of the signal. The quantitative and therefore exact
calculation of oxygen saturation does not only require constant path length but knowledge
of the path length. This problem was replied with the development of the time-resolved
spectroscopy, which can measure the blood volume and the oxygen saturation quantitatively
(Essenpreis et al 1993). This system uses time-correlate single photon counting method for
measuring the temporal function of the sample. The system measures the intensity of light in
a time domain and enables analysis of the data with the time domain photo diffusion equation
(Chance 1989). Due to the knowledge of the path length and constant light scattering, the
quantitative assessment of oxygen saturation can be carried out in patients (Ohmae et al 2006).
3.6. Orthogonal polarization spectral imaging
Direct visualization of the microcirculation was first achieved with capillaroscopy.
Capillaroscopy allows two-dimensional visualization of the capillary network in real time,
especially in the nailfold (Houtman et al 1986). The most widely used method to study
the morphology and function of skin capillaries is fluorescence videocapillaroscopy. In
combination with intravenous administration of fluorescent dyes, e.g. indocyanine green,
demarcated microvascular structures can be visualized, which otherwise are hardly visible
with ordinary static light microscopes. The skin melanin absorbs light strongly in the
visible spectrum. Therefore, visualization is limited mainly to the nailfold, and ordinary
videocapillaroscopy is difficult in pigmented skin (Yvonne-Tee et al 2006).
A further development of direct microcirculatory visualization is the orthogonal
polarization spectral (OPS) imaging. OPS is a new non-invasive method developed for the
assessment of the human microcirculation without using fluorescent dyes in clinical practice.
The instrument consists of a small endoscopic-like light guide attached to a light source with
filters. The examined tissue is illuminated with polarized light with a wavelength of 548 nm
permitting optimal imaging of the microcirculation, because of identical light absorption of
oxy- and deoxyhemoglobin at this wavelength. Within the tissue, light is scattered, depolarized
and reflected. Since the emitted light is primarily absorbed by hemoglobin, red cells can be
remarkably well observed in all vascular segments up to a tissue depth of approximately
300 µm(Harriset al 2000). The video images obtained are of high resolution and
microcirculatory parameters such as functional capillary density (number of capillaries with
flowing red blood cells mm
−1
), vessel type and diameters, type of red blood cell flux
(continuous, intermittent or absent) and blood flow velocity can be analyzed semiquantitatively.
OPS was validated in several animal experiments by comparison with fluorescence
intravital microscopy in particular in the hamster check pouch window preparation
demonstrating reasonably good agreement of measured microcirculatory parameters under
control conditions and after ischemic injury (Groner et al 1999).

R80 Topical Review
Figure 6. Visualization of the microcirculation in the sublingual mucosal area recorded with
the sidestream dark-field imaging technique. This picture shows diminished functional capillary
density in a patient on the heart–lung machine during cardiac surgery.
In humans OPS has been successfully applied to investigate the microcirculation of the
tongue, gingiva, vaginal mucosa, in burn wounds, skin, liver and the brain under different
conditions (Trzeciak et al 2007, Verdant and De Backer 2005, Lindeboom et al 2006,van
den Oever et al 2006, Milner et al 2005, Genzel-Boroviczeny et al 2004, Virgini-Magalhaes
et al 2006, Puhl et al 2005, Pennings et al 2006). Recent investigations suggest that OPS
imaging is suitable to monitor the efficiency of therapeutic maneuvers in order to improve
microcirculatory blood flow in patients with septic shock (De Backer et al 2006). In
these patients, the interrater variability (0.79–0.91) and intrarater variability (0.67–0.89) for
measurements of microcirculatory parameters in the sublingual region were reported to be
reasonably well (Boerma et al 2005).
A further development of this technique is the sidestream dark-field (SDF) imaging. SDF
imaging is based on slightly different principles as compared with OPS technology. Light-
emitting diodes arranged in a ring formation at the tip of the light guide emit green light with
a wavelength of 540 ± 50 nm, which directly illuminates the tissue microcirculation (Ince
2005). The illuminating light source is optically isolated from the emission light path in the
core of the light guide. SDF technology provides improved resolution and clarity of the images
compared to OPS imaging (figure 6)(Tureket al 2007).
Major limitations of OPS and SDF imaging are that only tissues with a thin epithelial layer
can be examined. Therefore, measurements are usually limited to the oral mucosa in adults and
additionally the skin in newborns. In intubated patients, the presence of blood and/or salive
may limit good visualization of the microcirculation. Uncontrolled pressure of the probe can
significantly decrease microvascular blood flow, and lateral movement of tissue or repeated
application of the device may prevent continuous investigation of a defined microvascular
region in the tissue. The current technical limit for blood flow velocity determination
is approximately 20 mm s
−1
and until today semiquantitative analysis of microcirculatory
parameters can be only performed ‘off-line’.
4. Conclusion
During the last decades, microcirculatory research has directed clinicians to a better
understanding of the pathophysiology of various diseases. New, non-invasive technical devices
for microcirculatory monitoring have been developed for application in clinical practice. And

Topical Review R81
new, emerging technologies will probably further elucidate microcirculatory hemodynamics
and tissue metabolism. Optical coherence tomography, for example, is based on low-coherence
interferometry that enables non-invasive, high-resolution, two- or three-dimensional, cross-
sectional imaging of microstructural morphology in biological tissue in situ (Gambichler
et al 2005). Also a deeper look in the oxygen metabolism of cells in situ at the mitochondrial
level promises better insights into microcirculatory function and oxygen supply mechanisms
(Mayevsky and Rogatsky 2007). Unfortunately, the interpretation of measurements in order
to predict the actual clinical status of a patient is limited mostly by substantial differences
in the regulation of microcirculatory blood flow between organs and by the wide variety
of measured parameters when comparing different monitoring systems. Finally, the cost
effectiveness of different monitoring devices designed to measure microcirculatory phenomena
has not been determined yet. In the critical care setting analysis of cost effectiveness
would require an exact determination of the probability of microcirculatory pathology in
a defined disease, e.g. septic shock. Furthermore, a positive outcome by improving the
microcirculation with a therapeutic intervention has to be convincingly demonstrated. In
this situation, routine monitoring of microcirculatory parameters would be assumed to be
cost effective if the product of the probability of microcirculatory disturbances and costs
related to a preventable adverse outcome would exceed the financial burden related to
routine monitoring of the microcirculation in that particular disease (Polk and Roizen 1990).
An ‘optimal’ microcirculatory monitoring technique for clinicians should evaluate ‘global’
microcirculatory function including microcirculatory blood flow, tissue oxygen delivery,
oxygen content and consumption, as well as tissue metabolism. Having a look on this
‘optimal’ microcirculatory technique, one has to consider two major problems. First of all,
as mentioned above, the microcirculation differs not only from one organ to another in its
anatomy and physiology, but also in the organs itself you will find different microcirculatory
physiologic and pathophysiologic alterations. For that reason, it is difficult to interpret data
from one organ bed and generalize these results to the whole body. The second main problem is
still the nescience of the interpretation of the data gathered with the present technical devices.
How much microcirculatory blood flow is enough for adequate tissue oxygen delivery? At
which tissue oxygen tension does a cell-breakdown occur? There are still no absolute levels
to describe and define microcirculatory function parameters. These problems make it very
difficult to develop new, better devices created for microcirculatory monitoring at the bedside.
Nevertheless, despite severe limitations, the monitoring of the microcirculation may, ahead of
time, alert physicians that tissue oxygen supply becomes compromised and leads to a better
understanding of basic pathophysiologic aspects related to disease.
Disclaimer
The authors have no financial interest in any of the products mentioned in this review.
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