Refractometry

Refractometry is a technique that detects the difference in refractive indices between two media, one of which contains the gas sample of interest (e.g. air and anaesthetic vapour), the other which is of otherwise identical constitution (e.g. air only). The refractive index of a medium is a measure of the ratio of velocity of light in a vacuum to the velocity of light in that medium. The molecules of the medium delay the transit of the light wave through it; the delay depends on the number of molecules and hence on the density of the medium; thus refractive index for a gas medium depends on its concentration, pressure and temperature. When a light beam passes through parallel slits whose width is of the same order of magnitude as the wavelength of the light, interference patterns are produced by light waves arriving in phase (bright fringe) and 180° out of phase (dark fringe) with each other. When two such sets of fringes are formed from light passing through gas samples with differential velocities, they are displaced relative to each other and the displacement between them can be measured as a difference in refractive indices.

The refractometer is included, not because it is frequently used clinically as a gas analyzer, but because it is an accurate standard of gas analysis against which others are compared. All gasses and volatile agents can be quantified using this method, since all possess the physical property of refractive index. Laboratory refractometers are used by vaporizer manufacturers to calibrate vaporizers. A diagram of the portable version of the refractometer (the Riken refractometer) is shown in Fig. 15.3. Light from a small light source is collimated into a beam, which is split into two parallel beams by prisms. One beam is passed through the reference chamber containing, say, air, and the other is passed through the measuring chamber containing air and another gas/vapour species, for example sevoflurane, whose concentration is to be measured. The beams are reflected back through the chambers to enhance the effect. Each beam is passed through a slit, producing interference fringes. As discussed above, the effective path lengths of the two beams differ and two sets of fringes are produced. These can be realigned by a knob controlling a vernier scale and the amount of realignment is taken as a measure of refractive index change from the reference sample, and thus of concentration of the gas under test.

Figure 15.3 The principles of the portable refractometer.

Infrared absorption spectroscopy

The interatomic bond between dissimilar atoms of a molecule absorbs radiation in the infrared (IR) range. Thus, molecules such as nitrous oxide, carbon dioxide, water vapour and the volatile agents absorb IR light, but oxygen, nitrogen and helium do not. Polyatomic molecules of different species absorb IR radiation across a range of IR wavelengths and there is frequently overlap between species. However, different polyatomic species absorb maximally at characteristic wavelengths within the IR bandwidth, which means that it is possible to identify the gas molecule as well as quantify the gas concentration. The amount of absorption (A) of any radiation by any substance is governed by the Beer-Lambert Law, which links A to the intensity of the incident radiation I_i, the non-absorbed, transmitted radiation I_t, the extinction coefficient ε, the path length L and the concentration C of the absorbing substance, where:

Since CO₂ analysis is important in anaesthesia, such devices are available in most anaesthetic locations in modern operating rooms. Fig. 15.4 shows the principle of the Luft or non-dispersive type of IR gas analyzer, which uses a differential technique to minimize error. An IR light source emits radiation at between 1 and 15 µm. A ‘chopper’ wheel, with small windows containing optical filters at the rim, rotates in front of the light source at 25–100 Hz to prevent the gas sample from overheating; additionally the filters allow the passage of IR light of a narrower bandwidth than the source itself, to match the wavelength of maximum absorption of the gas under study. This window might contain a 4.26 µm filter for CO₂, a 3.3 µm filter for halothane, isoflurane and enflurane, and a 4.5 µm filter for N₂O. The light passes through this to two chambers, a reference chamber and a sample chamber; the sample chamber contains the same background gas as the reference chamber, but additionally contains the gas under analysis. The light, which is not absorbed (the transmitted light) in each sample cell, is passed to a pair of air-filled detector chambers, separated by a diaphragm. Because different amounts of IR light are absorbed in the reference and sample chambers, different amounts of light are, therefore, transmitted to the detector chambers, heating the air in them differentially. Because the chopper wheel produces an oscillating IR source, the changing air pressure in the two detector chambers causes the diaphragm separating them to oscillate. The diaphragm usually forms half of a capacitor and is, therefore, a component of electrical circuitry, which amplifies and processes the signal, to produce an output signal, which is proportional to the gas concentration.

Figure 15.4 The principles of the infrared spectrophotometer.

There are a number of sources of error in IR spectroscopy. Although CO₂, N₂O and CO absorb IR light maximally at 4.3, 4.5 and 4.7 µm respectively, there is considerable overlap in the absorption spectra of these gasses, particularly in the 3–5 µm range. There is also the phenomenon of collision broadening, where the presence of one gas may broaden the IR absorption spectrum of another. This effect is not only caused by IR absorbing agents themselves (such as CO₂ and N₂O) on the absorption spectra of other IR absorbing agents, but by other non-IR absorbing carrier gasses, such as helium, argon or hydrogen. It has been reported, for example, that in a gas mixture containing 79% helium in oxygen, an IR analyzer under-reads CO₂ values.⁴ Desflurane,⁵ cyclopropane,⁶ acetone and alcohol⁷ all produce errors in IR spectroscopy of respiratory gasses. There are electronic correction factors in the analyzers to allow for the presence of N₂O in the gas mix containing CO₂. Since water is a strong absorber of IR across the bandwidths of interest, it must be eliminated in the sampling process, in order to avoid some inaccuracy. If one wavelength is used for all volatile agent detection, then that agent must be selected manually, or the device will be inaccurate, if either the wrong agent is selected or if there is a mixture of agents present. Some analyzers use the 10–13 µm bandwidth to detect volatile agents, where there is less chance for interference between absorption spectra of the volatile agent and other gasses.

Infrared spectroscopy detects numbers of molecules in a gas sample and thus constitutes a partial pressure analyzer. Therefore, error can be introduced if the pressure of the gas sample changes or if there is ambient pressure change, against which the device is calibrated. If the gas sample pressure changes due to back pressure from a ventilator, or if there is significantly low pressure in a sampling tube, the partial pressure of the gas being analyzed will change, without there being a real change in the fractional concentration of the gas. Similarly, if the device is calibrated at sea level and subsequently used at altitude, there will be an error in calculating gas concentration. Therefore, calibration should be carried out frequently and under the conditions of intended use, or the device should be configured to display partial pressure rather than concentration.

The 90–95% response time of IR analyzers to a step change, including transit time and rise time, should be under 150 ms. Water vapour blocking the sampling tube sometimes causes the response time to increase. To deal with water vapour, most devices have either a water trap or use sample tubing, which absorbs water vapour. IR capnographs are accurate to about 0.1% in a range of CO₂ up to 10%.

A variant of the IR analyzer described above utilizes a combination of IR and photo-acoustic spectroscopy. The gasses under test are drawn through a small measurement chamber that is irradiated by an IR source. This IR source is first ‘chopped’ by a rapidly rotating wheel that has three concentric bands of holes that cause the IR energy passing through the holes to pulsate at three different sampling rates. Each part of the now divided beam passes through a separate narrow band optical filter allowing through only that wavelength that corresponds to the maximum absorption of the gas/vapour to be tested (i.e. CO₂, N₂O and volatile agent). The components of the IR light, now differentiated both by sampling frequency and wavelength, are absorbed by the gas molecules, which are caused to vibrate at different frequencies, leading to a set of audible, fluctuating pressure waves. The families of audible pressure waves are detectable by microphone. Fig. 15.5 shows the Bruel and Kjaer model of the photo-acoustic IR spectroscope. Note that, since oxygen does not absorb IR light, this device has a separate sampling channel that uses magneto-acoustic spectroscopy to measure oxygen (see the section on Paramagnetic oxygen analyzers).

Figure 15.5 The components of the Bruel and Kjaer photo-acoustic and magneto-acoustic spectrophotometer.

The signal from all gaseous components is taken from the microphone and filtered electronically into the four components of O₂, CO₂, N₂O and volatile agent. Separation and, therefore, gas identification occurs by audio frequency identification and quantification of the gas occurs by audio-amplitude. Note that the device does not distinguish between different volatile agents. The advantages of photo-acoustic IR spectroscopy over conventional IR spectroscopy include stability, zero drift, reduced need for calibration over prolonged periods and fast response time.

A more recent design of infrared absorption spectroscopy does away with the moving parts of the chopper wheel. Light from a broad-spectrum infrared source passes through a cuvette containing sampled gas. The Dräger ILCA (Fig. 15.6) then uses a prism as a beam splitter. On the four sides of the multichannel detection chamber are appropriate infrared filters and sensors for the maximal absorption wavelengths of each gas. By using two such modules in series the gas bench can quantify CO₂, N₂O and the individual volatile agents in the gas mixture.

Figure 15.6 Working principles of Dräger ILCA gas analyzer.

Mass spectrometry

Mass spectrometry identifies gas molecules by a two-stage process; the molecules are first bombarded in a vacuum with electrons, which converts them into charged particles; the charged particles are then separated in a magnetic field according to the ratio of their mass to charge. The method can, therefore, identify and quantify all molecular species, providing there are no unexpected components in a gas mixture, such as the propellant in a bronchodilator or acetone in the expired breath of a diabetic. It is still considered to be the standard against which other gas monitoring techniques are compared, and is more frequently used as a laboratory technique than a clinical one. However, it was available for use as a time-shared, multi-site sampling device before the advent of cheaper IR spectroscopy. One example of a mass spectrometer is shown in Fig. 15.7A. Essentially it has three stages: the first stage is where the sampled gas is drawn into a low-pressure chamber. In the second stage, which is the main part of the device, the sample is drawn into an ionization chamber by an even lower pressure (about 1 mmHg), where the gas molecules are bombarded with electrons. Finally, they are drawn into a dispersion chamber, where they are influenced by a magnetic field. This third stage is where the deflected and separated beams of charged molecules are detected and the signal is processed and displayed.

Figure 15.7 A. Principles of a mass spectrometer. B. Components of a magnetic sector mass spectrometer. C. Components of the quadrupole mass spectrometer.

In the vacuum chamber of the second stage of the device, the bombardment of the molecules, by a transverse beam of electrons, results in ionization of the molecule, usually to a single positive charge. The ions of the different molecular species then accelerate towards an electrically negatively charged plate, the acceleration plate, and out through a small hole, the molecular leak, into the dispersion chamber. From here the path of the ions is influenced by a magnetic field. The ions are deflected according to mass; the lightest being deflected the most. The different species of gas are thus separated according to their mass: charge ratio. In the third stage, the ions reach the photovoltaic detectors, where the rate of arrival of the ions is proportional to the partial pressure of the gas. The signal is processed, amplified and displayed.

Two different types of mass spectrometer exist, depending on the way in which the magnetic field is produced in the second stage and the mechanism of ion detection in the third stage.

In the magnetic sector mass spectrometer (Fig. 15.7B) the magnetic field lies at right angles to the path of ions and is produced by a combination of electrical and fixed magnetic fields. By varying the voltage on the acceleration plate, the velocity of the ions entering the magnetic field can be changed. This allows the deflected ion beam to be directed across a single detector plate and different components to be detected in turn. It is, therefore, possible to separate components of the gas sample according to their mass : charge ratio.

More commonly seen these days, is the quadrupole mass spectrometer. This device is more compact and allows better discrimination between the ionic components from a gas mixture. The magnetic field is produced by a combination of DC electrical and AC radio frequency fields. As shown in Fig. 15.7C, the quadrupoles consist of four rods with opposite pairs electrically connected. By careful tuning of the radio frequency component of the magnetic field, only ions of a given mass : charge ratio proceed through the quadrupole to the detector, all other ions oscillating and colliding with the device. By a combination of changing the voltage on the acceleration plate and of judiciously tuning the magnetic field, a spectrum of mass : charge components can be detected and quantified. By scanning at 50 Hz, it is possible to produce a continuous record of gas concentrations.

The respiratory mass spectrometer is accurate, giving good gas identification and quantification, requiring only 20 ml min⁻¹ gas sampling rate, with a 100 ms response time. However, it does have some disadvantages. The second stage of the device operates under almost vacuum conditions; this requires a high-quality, continuously running pump. If the device itself is at some distance from the sampling site, as it used to be in the days of timesharing of a single device, significant delay time may be added to the response time. Water condensation can be avoided by heating the sampling tube, but the response time for water vapour may still be longer than for other components.

Some molecule types may lose two electrons rather than one in the ionization process, and, therefore, become doubly charged ions rather than single. They then behave within the magnetic field like an ion with half the mass, which leads to confusion in interpretation. Furthermore, the ionization process can lead to fragmentation of a molecule, so that a mass spectrum appears at the output rather than a single peak. However, this anomaly is useful to distinguish the gas components with the same molecular mass, which would otherwise be difficult to distinguish. These include N₂O and CO₂ (44 Daltons [Da]), or N₂ and CO (28 Da). N₂O is fragmented into NO, O_2, N₂, N and O, while CO₂ is fragmented into O₂, C₂, C and O. Rather than try to detect and distinguish both gasses at 44 Da, N₂O can be detected at a subordinate peak of 30 Da and CO₂ at 12 Da. Because the fragmentation is predictable, the amplitude of the subsidiary peak can be used as a measure of the parent peak and, therefore, of gas concentration. Using appropriate low pass and high pass filters, obfuscating peaks in the spectra can be removed.

Raman spectroscopy

Rayleigh scattering refers to the scattering of light by particles in its path of size up to one-tenth the wavelength of the light and occurs without any loss of energy or change of wavelength. Being wavelength dependent, this phenomenon gives us nature’s blue sky because of the increased scattering of blue light. However, a small fraction of the incident light, about 10⁻⁶, is scattered with a loss of energy and a change of wavelength characteristic of the molecule off which the light is being reflected; this is Raman scattering. Raman spectroscopy has been used in industry for years as a means of identifying solids, liquids and gasses, but has had to await the advent of powerful laser light sources and sensitive photocell detectors to be useful in a clinical setting for breath-by-breath analysis.

Fig. 15.8 shows a diagram of a Raman spectroscope, incorporating an Argon laser source of wavelength 485 nm, high reflectance mirrors to concentrate the laser beam, a gas sampling chamber, appropriate optics, a detection system, and microprocessor and display system. Table 15.1 shows characteristic wavelength or frequency changes, which identify different gasses. If plotted graphically, the amplitudes of the frequency shifted peaks are proportional to the gas concentrations.

Figure 15.8 The components of the Raman spectroscope.

Table 15.1 Frequency shifts in different gasses detectable by Raman spectroscopy (N₂O and CO₂ give two characteristic frequency shifts).

In Raman spectroscopy, each gas is independently analyzed, including CO₂, N₂O, volatile agents, O₂, N₂ and water vapour. This is done by first filtering the Argon laser light at 485 nm, then passing the light through a system of optics to collect and focus the scattered low level light, then through an electronic filtering system; the very low levels of light are then detected by a photomultiplier tube. The response time is 100 ms and the device is, therefore, capable of breath-by-breath analysis. The sample is not altered by the process and can, therefore, be returned to the breathing system, which is a significant advantage in low flow anaesthesia,⁸ although there is some overlap between gas species.⁹ However, the devices require a great deal of electrical power and are noisy.

The piezoelectric (Engstrom Emma) gas analyzer

A pair of piezoelectric crystals connected to an electrical power source is made to resonate, with a characteristic frequency difference. The frequency difference occurs because one of the crystals is coated in silicone oil, which absorbs a volatile anaesthetic agent to which it is exposed. This changes the natural frequency of the crystal, and the frequency difference between the crystals changes by an amount dependent on the amount of volatile agent absorbed. Manual identification of the agent is required, but the device is remarkably accurate and stable.¹⁰

The paramagnetic gas analyzer

Most gas molecules are repelled by a magnetic field and are, therefore, termed diamagnetic. Two gasses, oxygen and nitric oxide, are attracted into the field and are termed paramagnetic. This property enables oxygen concentrations to be analyzed and is due to the presence of unpaired electrons in the outer shell of an oxygen molecule, which is able to generate force in a magnetic field. A paramagnetic oxygen analyzer of the original type is shown in Fig. 15.9A. Two glass spheres, suspended between the poles of a magnetic field, are filled with nitrogen, a weakly diamagnetic gas. The glass spheres are arranged in a dumbbell shape, suspended by a thread, tensioned to keep the dumb-bell in the plane of the magnetic field. Zeroing should be carried out in the carrier gas destined to have oxygen added to it at a later stage. When a gas mixture containing oxygen is drawn through the analyzer, oxygen is attracted into the magnetic field, displacing the nitrogen filled spheres away from it. The detection system can either be a deflection measurement (Fig. 15.9A) or a null deflection type (Fig. 15.9B), which measures the current required in the circuitry to restore the pointer to its null point. These devices are accurate to within 0.1% O₂, but are adversely affected by pressurization, vibration, water vapour and high flow rates; there is also a slow response time of up to 1 min.

Figure 15.9 A. A simple paramagnetic oxygen analyzer. B. A null deflection paramagnetic oxygen analyzer.

With permission from Kenny GNC, Davis PDD (2003) Basic Physics and Measurement in Anaesthesia, 5th edn. London: Butterworth Heinemann.

An analyzer has been developed to overcome these disadvantages (Fig. 15.10). A sample of gas to be measured is drawn continuously through one of two capillary tubes at the same rate as a reference gas sample (usually air) is drawn through the other. A powerful electromagnet, producing a pulsed (on/off) magnetic field is placed over the junction of the two tubes. The magnetic field, when switched on, attracts the oxygen from both tubes in proportion to their concentrations. This causes an intermittent differential reduction in pressure upstream in the tubes that is detected and measured by a pressure transducer, and can be calibrated so it displays oxygen concentration. This is the basis of the Datex paramagnetic oxygen analyzer.

Figure 15.10 A modern paramagnetic oxygen analyzer.

The pulsed magnetic field may be replaced by an alternating one that has a frequency of 110 Hz. This produces differential oscillating pressures of 20–50 µbar in the capillary tubes. The oscillations are transduced into a sound signal, the amplitude of which is directly proportional to the O₂ concentration in the sample. This is magneto-acoustic spectroscopy and forms the oxygen analyzing device in the Bruel and Kjaer gas analyzer.

It has been reported that a gas mixture containing desflurane interferes with the accuracy of a paramagnetic oxygen analyzer.⁵

Dräger PATO

The paramagnetic property of oxygen can be exploited in different ways as the basis for oxygen concentration measurement. In the Dräger PATO O₂ sensor a very small measurement chamber lying between two large magnets has heating and temperature measuring elements on either side. The presence of oxygen in the gas mixture reduces the thermal conductivity of the gas between the magnets producing a rapid response oxygen analyzer (Fig. 15.11A and B).

Figure 15.11 A. Schematic of Dräger Pato paramagnetic oxygen analyzer. B. Dräger PATO O₂ sensor complete, and disassembled to show magnets and interposing sensor strip carrying gas pathway and measurement compartment (the diameter of the cylinders is about 5 cm).

Fuel cells and polarographic cells

These techniques are used for analyzing oxygen and are included together, since they are similar electrochemical techniques.

The Clarke polarographic electrode is shown in Fig. 15.12A. It consists of a cellophane covered platinum cathode and Ag/AgCl anode in a phosphate and KCl electrolyte, between which a potential difference of –0.6 V is applied by the battery as shown. At the cathode the following reaction takes place:

Figure 15.12 A. The Clarke polarographic electrode. B. The Galvanic fuel cell.

Electrons are provided by the cathode, to be consumed by the O₂ in the gas or blood sample. At the anode the following reactions take place:

The rate of diffusion of oxygen into the electrolyte solution is a rate-limited process, generating a current proportional to the pO₂ of the gas sample.

The problem with the polarographic electrode is the dependence on a battery and the reduction of any N₂O present at the cathode when it is contaminated by Ag⁺ ions. A fuel cell is a similar device, which consists of a gold cathode and a lead anode (Fig. 15.12B). The same reaction occurs at the cathode as in the polarographic electrode. A polarizing voltage is generated by the process and ‘fuel’ is consumed, such that the cell needs to be replaced at intervals of 6 months to a year. The fuel cell’s response time is slow, making them acceptable, but less suitable for breath-by-breath analysis than other methods.¹¹

Nitric oxide measurement

Nitric oxide (NO) is a pulmonary vasodilator and as such, has found its place in an intensive care environment. Since NO causes methaemoglobinaemia, it is important to be able to measure its concentration. Since nitrogen dioxide (NO₂) is a degradation product of NO and causes pulmonary oedema, it too must be detectable.

All industrial analyzers developed for NO analysis are designed for use in environmental control and, therefore, for gas mixtures containing 21% or less of oxygen. Medical gas mixtures frequently contain more oxygen than this and the accuracy of devices available is not guaranteed. The two methods available are the chemoluminescent and the electrochemical methods.¹²

The chemoluminescence analyzer is based on the reaction of NO with ozone (O₃) in a chamber to produce energized NO₂ molecules, which emit photons of light in the range 590–2600 nm. The amount of light emitted is proportional to the original concentration of NO. In a parallel system there is a catalytic converter before the reaction chamber to reduce NO₂ in the sample to NO. The total oxides of nitrogen are then measured as NO and the difference in the two readings is the NO₂ concentration.

The electrochemical analyzer generates a potential difference proportional to the NO₂ concentration. This is due to the migration of ions generated from the reaction between NO or NO₂ and the electrolyte. A separate cell is used for NO and NO₂.

pH electrode

A pH unit, which is the measure of hydrogen ion activity in a liquid, is defined as:

In words, this can be described as ‘the negative logarithm, to base ten, of the hydrogen ion concentration’.

The physical principle, on which the pH electrode is based, depends on the fact that when a membrane separates two solutions of different [H⁺], a potential difference exists across the membrane. In a pH electrode, the membrane is made of glass and the development of a potential difference between the two solutions is thought to be due to the migration of H⁺ into the glass matrix. If one solution consists of a standard [H⁺], the pH of the other solution can be estimated by measurement of the potential difference between them. The glass membrane used is selectively permeable to H⁺. No current flows in this device, which therefore does not wear out, in contrast to the Clarke electrode, in which current does flow and which does need periodic replacement.

The pH measurement system is shown diagrammatically in Fig. 15.14 and consists of two half electrochemical cells. It has in one half an Ag/AgCl electrode and in the other a Hg/HgCl₂ (calomel) electrode. Each electrode maintains a fixed electrical potential. The Ag/AgCl electrode is surrounded by a buffer solution of known pH, surrounded by the pH sensitive glass. Outside the glass membrane is the test solution, usually blood, whose pH is to be measured. It is the potential difference across the glass, between these two solutions, which is variable. The blood or other solution is separated from the calomel electrode by a porous plug and a potassium chloride salt bridge to minimize KCl diffusion. The potential difference across the system is about 60 mV per unit of pH change at 37°C. The internal electrical resistance is high and in order to maximize the device’s accuracy, a voltmeter (or other measuring device) of very high internal resistance must be used to minimize current drawn from the system.

Figure 15.14 The pH electrode.

The pH electrode actually responds to hydrogen ion activity rather than hydrogen ion concentration. These two variables coincide at infinite dilution of solution, but might otherwise differ, because of molecular interaction between ionic species. The electrode is, therefore, calibrated against standard buffer solutions of pH 6.841 and 7.383. The glass of which the pH sensitive electrode is made consists of 72% SiO₂, 22% Na₂O and 6% CaO.

The Severinghaus pCO₂ electrode

There is a dynamic equilibrium between H⁺ and CO₂. The Henderson-Hasselbalch equation describes this relationship:

where α is the Ostwald solubility coefficient, 0.003 mmol L⁻¹ mm Hg⁻¹ at 37°C. pK_a, the association constant of the chemical reaction concerned is 6.1. Both α and pK_a vary with pH and thus lend the calculation some inaccuracy. It was Astrup who noticed that there was a linear relationship between pH and pCO₂.

The pCO₂ electrode is essentially a pH electrode with a difference. The sensitive glass membrane is covered with another membrane, which is selectively permeable to CO₂. In the original design, a layer of water was trapped between the two membranes, allowing the reaction described above to occur. A later modification had salt and bicarbonate solution in this space (Fig. 15.15). A pCO₂ electrode, therefore, allows a pCO₂ change to generate a pH change, which is measured by the electrode.

Figure 15.15 A pCO₂ electrode.

Derived variables from a blood gas machine

Buffer base is the sum of all bases in the blood capable of buffering pH changes. These include HCO₃⁻, Hb, PO₄³⁻ and the anionic parts of protein. A respiratory driven change in status, marked by, for example, the addition of CO₂, results in the formation of carbonic acid, which produces H⁺ and HCO₃⁻. The hydrogen ions are buffered by the total buffer base, namely: HCO₃⁻ + Hb⁻ + PO₄³⁻ + Protein⁻.

The buffer base can be considered to be a limitless source of buffering capacity, unless the haemoglobin component is significantly reduced. The base excess, an indication of metabolic rather than respiratory acid–base disturbance, is defined as the amount of titratable acid required to correct a blood sample to a pH of 7.40 at a pCO₂ of 5.3 kPa and at 37°C.

Standard bicarbonate is the bicarbonate in a fully oxygenated blood sample at a pCO₂ of 5.3 kPa and at 37°C.

Errors in blood gas measurement include:

• air bubbles in the sample

• excess heparin

• gas storage in container plastic walls

• failure to store samples at low temperature

• metabolism in the sample

• signal processing errors

• at the O₂ electrode:

 O₂ consumption in a Clark electrode

 material used in electrode membrane

 polarizing voltage

• at the pH electrode:

 protein deposition on the electrode.

Temperature and blood gas analysis

Blood gas analyzers measure blood gas variables at 37°C. A fall in body temperature means that CO₂ is more soluble in blood and that CO₂ production slows, leading to a fall in pCO₂. This shifts the O₂ dissociation curve to the left, increasing the O₂ content. However, O₂ solubility also rises, decreasing pO₂ and measured O₂ content. pH increases with a fall in temperature. Hypothermia itself does not alter the blood gas values at 37°C and arterial blood taken at lower temperatures should be corrected to 37°C before clinical decisions are made on the results.¹³

Other ion-selective electrodes

By carefully selecting the proportions of constituents in the glass, glass ion-selective electrodes can be constructed for detection of other ions. For example, a glass electrode for measuring Na⁺ concentrations consists of 71% SiO₂, 18% Al₂O₃ and 11% Na₂O. The K⁺-sensitive glass electrode consists of 69% SiO₂, 4% Al₂O₃ and 27% Na₂O.

Transcutaneous blood gas analyzers

pO₂ can be measured transcutaneously by applying a Clarke polarographic electrode to the skin. If the skin is ‘arterialized’ by heating it, a reasonable value for pO₂ can be obtained. A modified pCO₂ electrode is used to measure transcutaneous pCO₂. A common O₂/CO₂ permeable membrane and electrolyte solution is used, ensuring a common pH for both measurements. Actual measured values do not correlate well with arterial values and should not replace formal respiratory gas analysis,¹⁴ but trends are clinically useful and these devices are used particularly in neonatology and when patients are ventilated outside the intensive care setting where arterial blood gas sampling is problematic. Care must be taken to avoid burns due to the heated electrodes.

Intravascular blood gas analyzers

These are microminiaturized versions of the devices described above.

The co-oximeter

The co-oximeter is usually considered alongside the blood gas analyzer. The co-oximeter measures the fractional oxygen saturation of haemoglobin, taking account of the numerous different species of haemoglobin present in the blood, such as reduced Hb (HHb), oxygenated Hb (HbO), methaemaoglobin (MetHb), sulphaemoglobin and caorboxyhaemoglobin (COHb). The co-oximeter is a spectroscope, which, therefore, needs four or five different wavelengths to distinguish these different species. The Beer-Lambert law relating to light absorption is more closely adhered to than the pulse oximeter, which measures functional oxygen saturation, using only two wavelengths of light, and assumes the sample only contains reduced and oxygenated haemoglobin.

A small sample, 200–400 µl, of heparinized blood is injected into a cuvette of known path length. The red blood cells are disrupted by ultrasound or a soap solution, releasing free haemoglobin. Conventional spectroscopy then takes place on the sample at up to 17 wavelengths, which are produced either by a filter wheel with multiple narrow band filters, or by scanning with a diffraction grating or a prism. Fig. 15.16 shows the spectrum covered by a co-oximeter as well as that covered by a pulse oximeter. The result is usually available within 30 s and includes SO₂ %, total haemoglobin (Hb) concentration, oxygen content, methaemoglobin (MetHb) %, sulphaemoglobin %, carboxyhaemoglobin (COHb) %. Some devices are capable of distinguishing foetal haemoglobin (HbF). As Fig. 15.16 shows, the pulse oximeter and the co-oximeter operate in different spectra and, therefore, strictly speaking, SO₂ and SpO₂ should not be compared.

Figure 15.16 Absorption spectra for reduced and oxygenated haemoglobin showing the range of operation of a co-oximeter and of a pulse oximeter.