Although the cause of hydrocephalus is much more complex than an abnormality of the CSF circulation, shunt surgery is still considered the treatment of choice in most types of communicating hydrocephalus. The role of a shunt is simple: to drain excess CSF from the fluid spaces of the cranial-spinal compartment into an extracranial compartment. Hydrocephalus is caused by different etiological factors, but the common final pathway creates a vicious circle of altered CSF circulation, cerebral blood flow, and metabolism that affects brain homeostasis. In communicating hydrocephalus, draining of CSF appears to be the simplest way of interrupting this circle by improving the CSF circulation. Similarly, in noncommunicating hydrocephalus, third ventriculostomy improves the CSF circulation by bypassing the block to the anatomical circulatory pathway. If other ways of treatment exist, they remain to be demonstrated.

It is believed that an ideal shunt should contribute to restoration of the normal circulation of CSF and regional cerebral blood flow in areas adjacent to ventricular surface [1], prevent excessive build-up of intracranial pressure, and encourage restitution of the cerebral mantle, comprising both gray and white matter. As shunting is a purely mechanistic treatment, which moreover dramatically changes the biomechanics of a patient’s pressure–volume compensation, the CSF dynamics should be carefully examined before the shunt is implanted [2]. A shunt should correct the disturbances of CSF circulation by reducing the resistance to CSF outflow to about 5–10 mmHg/ml/min – the typical hydrodynamic resistance in normal brain [3]. Therefore, if a patient presenting with hydrocephalus-like symptoms has normal intracranial pressure, and a normal resistance to CSF outflow (6–10 mmHg/ml/min), a shunt theoretically cannot help because the resistance cannot be reduced further without the risk of overdrainage.

The ability to obtain variables that describe CSF dynamics before shunting allows clinicians to more categorically address some key questions in the clinical decision-making process:

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  Scope for improvement with CSF shunting: shunts work by providing an alternative (normal resistance) pathway for CSF outflow.

  
  
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  To determine which shunt system is the best to use for a patient and in the case of an adjustable valve, to choose initial shunt setting.

  
  
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  In case of shunt failure, later on, a pre-shunting infusion test provides a baseline for post-shunt comparison.

For all the above points, knowledge of the hydrodynamic properties of hydrocephalus shunts is paramount.

Shunt failure occurs at a rate of approximately 20–30% in the first year, and 3–5% per annum subsequently [4]. Clinical symptoms of shunt malfunction can be diverse for different subjects. Deterioration can be gradual and subtle or sudden and dramatic; moreover, for patients with comorbidities it can be nonspecific. In such cases it can be difficult to justify revision without quantitative evidence and comparative analysis. Similarly, in patients presenting with persistent symptoms post-shunting, testing of CSF dynamics is a useful method of excluding revision, if relative normalization of CSF dynamics has been demonstrated [5]. Assessment of shunt functioning in vivo is impossible without detailed knowledge of the shunt’s hydrodynamic properties.

The mathematical model of CSF pressure–volume compensation, introduced by A. Marmarou [6] and modified in later studies [7,8], provides a theoretical basis for the differential diagnosis of hydrocephalus. Shunt testing in vivo in hydrocephalus using physiological measurements requires the model to be extended by a branch representing a shunt (Figure 17.1 – part of the circuit, to the right from vertical line). It is important to notice that the reference pressure here is not the same as in the branch representing the patient’s physiological CSF reabsorption (i.e. sagittal sinus pressure). The hydrocephalus shunt bypasses the impaired physiological pathways of CSF circulation and creates a new nonphysiological connection between the CSF pools and the site of drainage (abdomen, right atrium, pleural space, etc.). Therefore, the reference pressure (outlet pressure) is that of the implantation site. Moreover, it could be grossly modified by the patient’s body position: decreased (in vertical position) by still controversial “siphoning pressure” or almost atmospheric in horizontal position. Theoretically it represents the pressure exerted by a water column in a tubing connecting ventricles to receiving extracranial compartment and reflects the distance separating them. Therefore, siphoning is greater in ventriculoabdominal than ventriculoatrial shunts [9].

Figure 17.1 (a) Electrical model of CSF circulation and implanted shunt. All parameters are strongly nonlinear. The pressure–volume curve (b) has a linear-exponential character. Resistance to CSF outflow (c) is infinitely high when ICP is below sagittal sinus pressure level and constant (6–10 mmHg/[ml/min]) when it is greater than sagittal sinus pressure. On top of this nonlinearity, the assumption that sagittal pressure is constant may not be valid in some clinical situations. Finally, nonlinearity of shunt performance may be substantial (d) as in the case of the Orbis-Sigma valve. The combination of these three nonlinear elements (pressure–volume curve, resistance to CSF outflow, and shunt flow–pressure curve) may make analysis of stability of ICP in shunted patients quite complex.

The resistance of most passive shunts is nonlinear. A diode with positive polarization voltage represents the valve: this can be of fixed or magnetically adjustable value. Hydrodynamic parameters of the shunt, i.e. its pressure–flow performance, hydrodynamic resistance to flow when the shunt is open, and opening/closing pressure, will be discussed in detail in this chapter. Knowledge of these parameters is important when choosing which valve to implant, and in adjustable valves choosing an appropriate setting; it is also important for confirming shunt function in vivo when adverse clinical symptoms occur after shunting.

Hardware

From the first documented cases of shunting in 1896 by Mikulicz, the attempts to develop a reliable internal CSF shunt turned into an avalanche in the second half of the twentieth century [10]. The recent technological improvements have been substantial: from the first rubber ventriculojugular valves invented by Nulsen and Spitz in 1952 to contemporary high-tech adjustable devices. New materials prevent bacterial adhesion and colonization, novel constructions help to avoid overdrainage. A wide variety of shunt products (over 80 generic types, having numerous subtypes and performance levels) are on the market, with little systematic knowledge available to help the practicing surgeon compare their comparative cost-effectiveness. Similarly, there is very little known about whether a specific type of shunt is better suited to correct an individual pattern of disturbed CSF circulation.

Classical differential-pressure valves: most contemporary valves are differential-pressure passive constructions. This means that the drainage of CSF is regulated by pressure difference between the inlet and the outlet of the valve. Consequently these valves control the pressure in the CSF compartment from where the fluid is drained. The opening pressure of the valve is the determining parameter. And this opening pressure can be fixed or (in the new generation devices) adjustable magnetically after implantation. There are only a few constructions that are designed to regulate the flow through the shunt in a range of differential pressures across the valve. Classification, construction principles, trade names, and hydrodynamic resistances of the most popular valves are listed in Table 17.1.

Table 17.1 The Most common types of shunts and their hydrodynamic resistances (without distal drain)

cd, classic differential fixed pressure shunt; sp, membrane or other (SiphonGuard) siphon preventing mechanism included; * discontinued.

In a horizontal body position, the mathematical equation describing the drainage through the shunt is relatively simple to solve, providing the pressure–flow performance curve of the shunt is known. In the vertical position, everything is disturbed by the gravitation-induced pressure gradient in the long tubing (average 80 cm long), connecting the inflow and outflow sides of the shunt. This may accelerate CSF drainage and consequently an additional decrease in CSF pressure in the vertical position (on average by 10 mmHg), potentially causing symptoms related to postural overdrainage: headaches, subdural collections, or even hematoma. Another complication related to overdrainage is the “slit ventricle syndrome” [11] affecting mostly young patients. This condition, characterized by completely collapsed ventricles and elevated (30–40 mmHg) intraparenchymal ICP and responsible for severe symptoms of intracranial hypertension, is difficult to correct and in some cases even life-threatening.

Devices intended to attenuate the influence of postural changes on the drainage (antisiphon devices) can be incorporated in the valve body or can be added in line with the valve.

Lumboperitoneal shunts may produce similar posture-related problems. The difference is that the hydrostatic pressure does not produce pulling but rather pushing force in the vertical body position. Therefore, designated lumboperitoneal shunts have usually greater hydrodynamic resistance. If low hydrodynamic resistance valves are used in this arrangement, adverse effects such as tonsilar herniation are observed frequently [12].

The hydrodynamic performance of shunts is complex and determined predominantly by the design of the valve. Most modern constructions use either a ball-on-spring or silicone membrane to close the main CSF drainage aperture. For a summary of different constructions, see Figure 17.2a. Miter, distal/proximal slit, and cruciform valves are more common in the older generation of shunts. Some specific designs such as flow-regulating or referential pressure valves (so-called “Beverly Shunt,” now discontinued), posture-compensating gravitational valves, etc., cannot readily be classified along traditional lines.

Figure 17.2 The most popular constructions of shunts and components. (a) Membrane valve (upper) and ball-on-spring design (bottom). (b) Adjustable valve: Sophy (left) and Codman Hakim adjustable (right). All adjustable shunts have ball-on-spring valves. (c) Siphon-control devices: membrane (upper) and gravitational (bottom).

In the late 1980s the concept of adjustable valves was developed in the laboratories of major shunt manufacturers. Constructions used by then, and continued only with modest modifications until today, were ball-on-spring (Figure 17.2b). Adjustability is achieved by changing the pre-load of the spring. Moving a special rotor with an externally applied strong magnetic field changes the pre-load, producing matching different opening pressures of the valve. The range of adjustments usually varies within limits from 1 to 15 mmHg, in a variable number of steps from 3 to 20 (Table 17.2). In most constructions the hydrodynamic resistance of the valve remains constant when the operating pressure is changed. Once the valve is opened, it represents fixed low resistance to the flow.

Table 17.2 Comparison of most popular adjustable valves

Adjustability theoretically facilitates fine-tuning of the valve’s performance in vivo and better control of hydrocephalus, avoiding the need for revision. There are drawbacks: with the exception of a few of the newest constructions, adjustable devices can be accidentally affected by the external magnetic field.

Siphon-preventing devices can be divided into membrane, gravitational, and flow-regulating constructions (Figure 17.2c). Membrane constructions theoretically cut the flow through the shunt system when pulling (distal tube) pressure exceeds pushing pressure (ventricular) with a ratio 20:1 or so. They need a reference atmospheric pressure acting on the membrane, theoretically transmitted through the skin. However, in fat patients or in cases where scarring hardens skin above the membrane, these types of devices may become unreliable. Furthermore, increased membrane pressure may stop the flow through the shunt altogether. Gravitational devices (such as Shunt Assistant) increase opening pressure in upright body position by adding extra weight acting on the ball controlling flow in the device in-line with a shunt system.

Flow-regulating devices (such as SiphonGuard or Orbis-Sigma Valve) increase hydrodynamic resistance of the device when flow increases. This theoretically is able to limit overdrainage related both to body posture and to nocturnal vasocycling [13].

Laboratory shunt testing

Over the years, a number of independent laboratories, usually supported by academic institutions, have set the standard for shunt testing in vivo. Testing of shunts in Europe originated in the laboratory of Dr. A. Ashoff in Heidelberg over two decades ago [14]. The UK Shunt Evaluation Laboratory (Cambridge, UK) was established almost 20 years ago thanks to a grant from the Department of Health. Over this period, 26 shunts have been evaluated according to the ISO 7197 standard. Under the initial grant (1993–1998) all shunts in use in the UK were systematically evaluated with “blue reports” printed by the Medical Devices Agency. New devices were tested as they appeared on the market (or as prototypes), and these results have been published in academic journals.

The Shunt testing rig is depicted in Figure 17.3a. Measurement is controlled by a standard IBM compatible personal computer that reads and zeroes the balance periodically (every 15 s) to calculate the drainage rate. In this way the weight of the outflowing fluid is measured incrementally, which cancels the influence of fluid vaporization from the outlet . The computer analyzes the pressure waveform from the pressure transducer and controls the rate of the infusion pump. The effect of changes in atmospheric pressure is compensated for by using the reference barometer.

Figure 17.3 Principles and results of shunt testing in vitro: (a) Shunt testing rig. Computer controls measurement process. Shunt may be perfused with controlled rate (from infusion pump) or with decreasing pressure (from cylinder). (b) Flow–pressure (flow profile increasing and decreasing in a triangular manner) and (c) pressure–flow (pressure decreases as water head from cylinder decreases) tests. (d) Pressure–flow performance curves for fixed pressure shunts at three different performance levels (low, medium, high).

The shunt and pressure transducer are placed on the same level. The water column in the fluid container (H), the degree of the shunt submersion, and the level of the outlet tubing (O) may be changed according to the test protocol.

The shunt is tested under two different regimes:

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  Flow–pressure test: Controlled fluid flow results in a differential pressure across the shunt (Figure 17.3b). Flow is the independent variable and the resulting differential pressure is the dependent variable. The computer controls flow by changing the rate of the infusion pump.

  
  
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  Pressure–flow test: Controlled differential pressure across the shunt causes fluid to flow (Figure 17.3c). Pressure is the independent variable and flow is the dependent variable. Pressure is controlled by means of a water column (height H) in the fluid container (with stopcock A open horizontally). This pressure decreases during the test as fluid flows out of the container.

The pressure–flow and flow–pressure relationships are plotted during every test (Figure 17.3d). The following parameters are calculated:

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  Closing pressure: differential pressure below which flow through the shunt ceases. The closing pressure is measured as the intercept of the regression line with the x-axis, drawn between pressure (independent variable) and flow (dependent variable) for flow from 0.2 to 0.05 ml/min.

  
  
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  Hydrodynamic resistance: the change in pressure divided by the change in flow decreasing from about 1.2 to 0.3 ml/min. The resistance is measured as a linear regression gradient between pressure and flow. It is the resistance of the permanently opened shunt and is expressed in mmHg/(ml/min). In adjustable valves the resistance is commonly confused with closing pressure.

  
  
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  Differential pressure for physiological flow (0.3 ml/min) – so-called “operating pressure.”

Hydrodynamic properties of contemporary shunts

Eighteen nonprogrammable and eight programmable valves reveal the common hydrodynamic properties of contemporary shunts.

For different construction mechanisms, pressure–flow performance curves may have different shapes from completely linear (above opening pressure) to absolutely nonlinear (e.g. flow-regulating valves like Orbis-Sigma or Diamond). In some valves wide hysteresis of the performance curve can be seen, suggesting that measured differential pressure is dependent on flow through the shunt; silicone valves show particularly significant hysteresis whereas ball-on-spring valves are less susceptible. Ball-on-spring valves show usually better convergence of pressure–flow performance characteristics than membrane valves. Some constructions such as distal-slit valves may change their performance dramatically if distal arrangements change (wet or dry end, touching tissue at the outlet, or outlet being suspended in fluid).

Hydrodynamic resistance is calculated for pressures greater than valve opening pressure and is a well-defined parameter for valves with relatively linear pressure–flow performance curve; it cannot be evaluated for the Orbis-Sigma or Diamond valves (for these valves, stabilizing flow, hydrodynamic resistance is very high; theoretically infinite). In the literature, hydrodynamic resistance is commonly confused with operating pressure. The majority of the contemporary classic-differential shunts show low resistance to flow (as low as 1.05 mmHg/ml/min, which is substantially lower than physiological resistance to CSF drainage, measured at 6–10 mmHg/[ml/min] in normal subjects [15]). Low resistance is likely to result in overdrainage of CSF. Exceptions are the Medtronic Lumbo-Peritoneal Shunt, Codman Uni-Shunt, Sinu Shunt, and to some extent Holter Valve, of which the latter two have been discontinued. The resistance of the Uni-Shunt, however, may be strongly affected by conditions for flow at the distal end (e.g. in a peritoneal cavity).

Any repetitive variations of proximal pressure have a tendency to decrease the nominal operating pressure of shunts with unidirectional valves. This may lead to overdrainage in situations with regular vasogenic ICP waves or high respiratory fluctuations, particularly often seen in lumbopleural implantation. For this reason, lumbopleural shunts should be chosen from those with greater hydrodynamic resistance.

Long distal catheters increase resistance of the majority of classic-differential valves towards normal physiological values. It is important to remember that the resistance of a catheter is an inverse of the fourth power of its inner diameter and it is directly proportional to its length (Poisseuille’s law). Therefore, a 1 m long catheter with a 1 mm inner diameter has a resistance of around 5 mmHg/(ml/min), while a similar length catheter of 1.2 mm inner diameter has a resistance of around 2.5 mmHg/(ml/min). A useful formula to calculate expected hydrodynamic resistance of the distal drain (it is always better to measure it – the resistance may be dependent on the materials used and the finish of the inner surface) is:

where D is an inner diameter expressed in millimeters and L is length expressed as a fraction of one meter (the above formula is valid for density and viscosity of CSF close to parameters of water).

By comparison, the resistance of the ventricular catheter is not greater than 1 mmHg/(ml/min). The number of patent holes in ventricular catheters does not usually change the resistance of the tubing as the resistance is mainly affected by the tube itself.

All valves with membrane siphon-preventing devices are sensitive to external pressure. External pressure exerted by tense skin or scar on the skin increases operating pressure of the valve. Increased external pressure (cap, headband) may close CSF drainage completely. External compression of the membrane is used in shunt testing in vivo to reveal patency of ventricular drain. All constructions without membrane siphon-preventing devices are not sensitive to external pressure up to 50 mmHg.

Negative outlet pressure decreases operating pressure by the same value in all valves without a siphon-preventing mechanism. When the resistance of the shunt system is low (4–6 mmHg/[ml/min]), negative outlet pressure of –15 mmHg may accelerate the drainage rate to a nonphysiological value of 2–4 ml/min. Overdrainage may also occur when a low resistance valve is subjected to repetitive cycling of proximal pressure (exceptions are Orbis-Sigma, Diamond Valve, and valves fitted with Codman SiphonGuard). Another rarely mentioned cause of overdrainage is “pumping” of proximal reservoir of the shunt [16], which is commonly performed in emergency departments when shunt dysfunction is suspected.

All adjustable valves can be reset in vivo by applying an external magnetic field. Most settings cover a range of operating pressures from 0 to 20 cmH₂O (0–15 mmHg). The number of steps varies from 5 to 20 (Table 17.2). In almost all valves, the levels are equally spaced, except for the Codman Certas Valve (Figure 17.4). In all valves except the Codman Hakim Programmable Valve, verification of the setting may be conveniently performed without the need for radiography using an external compass placed over the valve. Both measurement and adjustability may be affected if the valve rotates under the skin.

Figure 17.4 Shunt operating pressures assessed for different types of adjustable valves. (a) First adjustable valve Sophy (SU3), three levels adjusted magnetically. (b) Next generation Hakim-adjustable valve with 18 possible settings. This construction was much more precise, but a disadvantage was that the setting had to be confirmed with X-ray. Both constructions could be reset by a relatively weak magnetic field (20–40 milliteslas).

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