ECG - input impedance and noise

In ECG standards, clauses such as 201.5.3 cc) of IEC 60601-2-25 require that test signals are accurate to ±1%. Although not explicitly stated in the standard, it’s obvious that this includes noise. If for example a test requires the ECG under test to reproduce a test signal accurately within ±5%, it would make no sense to perform the test if the test signal also had say ±20% noise.

Some engineers however might consider environmental (common mode) noise as a separate issue. ECGs are required to work in noisy environments so they should have the ability to reject common mode noise. Hence testing in a normal “noisy” environment is representative of the real world. This is wrong for two related reasons: first the ECG is anyhow tested for the ability to reject noise including common mode noise; second is that in order to test objectively, we need to start with a low noise environment and then add test signals including noise in a known and accurate way.

As such, while there are no instructions, notes, requirements or methods in the standard to minimise noise, such actions are implicit in the requirement to apply test signals with an accuracy of ±1%. In particular, most test signals are in the order of 1mV, which means noise of just 10µV is significant. Normal methods to minimise noise include using a ground plane under the ECG, cables and test equipment, and connecting the ECG equipment earth (PE or FE) and the test circuit ground to the ground plate.

The input impedance test is by far the most sensitive to noise, and sometimes the normal procedures are not enough. The reason for the high sensitivity to noise is the large imbalance impedance. To understand why this is so, it is useful to review the article on CMRR testing, which explains how CMRR is really a function of leakage currents flowing through the imbalance impedance. It follows that the size of the imbalance impedance directly impacts the size of the CMRR noise on the ECG.

For CMRR testing, the imbalance is 51kΩ, while for the input impedance test the imbalance is 620kΩ, some 12 times larger. This means that the input impedance test is 12 times more sensitive to noise than the CMRR test.

We can do some ball park calculations to illustrate the point. If say in an CMRR test the ECG records a 3mm indication; in voltage terms this is 0.3mVpp from a 10Vrms common mode voltage.

In the input impedance test, the typical test voltage is 3.2mVpp (80% of 40mm channel width @ 10mm/mV), so a 1% error is roughly 0.03mVpp, which is 1/10 of the above CMRR result. Since the set up is 12 times more sensitive to noise, it means the common mode voltage in the input impedance test needs to be less than:

Vcm = 10Vrms / 10 / 12 = 0.083Vrms = 83mVrms

A floating circuit with cables can easily pick up 2~10Vrms of common mode voltage from the environment, and people near a test site tend to make this worse. So, to get this down to 83mVrms may need some special shielding.

Additional measures may include adding a shield above the test set up (in particular for ECG cables), having the operator touch the ground plate during the test, and/or keeping ac power cables away from the test area as far as possible. The thickness of the shielding materials also helps: being very thin aluminium foil is sometimes not much help, but a 1mm thick aluminium plate usually works well.

Although not specified in IEC 60601-2-25 and IEC 60601-2-27, it is also possible to turn on the AC filter to remove 50 or 60Hz noise. The AC filter will reduce the common mode noise, should have no effect at 0.67Hz and may slightly reduce the signal at 30Hz. For example, you may need 3.5mVpp to get 32mm on the ECG due to the AC filter. Since the input impedance test is ratiometric, as long as the filter is on during the whole test the test result will still be valid.

Finally it is worth to note that some environments are naturally quiet while others are incredibly noisy. The input impedance test condition with the 620kΩ in circuit is a good worst case condition to check out the noise levels at different test sites. It can be useful to select a quiet site for all performance tests.

IEC 60601-2-25 Clause Input Impedance

The input impedance test is fairly simple in concept but can be a challenge in practice. This article explains the concept, briefly reviews the test, gives typical results for ECGs and discusses some testing issues. 

What is input impedance? 

Measurement of voltage generally requires loading of the circuit in some way. This loading is caused by the input impedance of the meter or amplifier making the measurement. 

Modern multimeters typically use 10MΩ input for dc measurements and 1MΩ for ac measurements. This high impedance is usually has negligible effect, but if the input impedance is similar to the circuit impedance significant errors can result. For example, in the circuit shown, the real voltage at Va should be exactly 1.000Vdc, but a meter with 10MΩ input impedance will cause the voltage to fall by about 1.2%, due to the circuit resistance of 240kΩ.   

Input impedance can be derived (measured) from the indicated voltage if the circuit is known and resistances are high enough to to make a significant difference relative to noise and resolution of the measurement system. For example in Figure 1, if the supply voltage, Rs and Ra are known, it is possible to work back from the displayed value (0.9881) and calculate an input impedance of 10MΩ. 

The test

The test for ECGs is now largely harmonised in all the IEC and ANSI/AAMI standards, and uses a test impedance of 620kΩ in parallel with 4.7nF. Although the requirement is written with a limit of 2.5MΩ minimum input impedance, the actual test only requires the test engineer to confirm if the voltage has dropped by 20% or less, relative a the value shown without the test impedance.  

ECGs are ac based measurements so the test is usually performed with an sine wave input signal. Input impedance can also change with frequency, so the IEC standards perform the test at two points: 0.67Hz and 40Hz. Input impedance test is also performed with ±300mV offset, and repeated for each lead electrode. That makes a total of 2 frequencies x 4 conditions (reference value + open + 2 offsets) x 9 electrodes = 72 test conditions for a 12 Lead ECG.  

Typical results

Most ECGs have the following typical results:   

  • no measurable reduction at 0.67Hz
  • mild to significant reduction at 40Hz, sometimes close to the 20% limit
  • not affected by ±300mV dc offset 

Issues, experience with the test set up and measurement

Although up to 72 measurements can be required (for a 12 Lead ECG), in practice it is reasonable to reduce the number of tests on the basis that selected tests are representative. For example, Lead I, Lead III, V1 could be comprehensively tested, while Lead II, V1 and V5 could be covered by spot checks at 40Hz only, without ±300mV.

In patient monitors with optional 3, 5, and 10 lead cables, it is normal to test the 10 lead cable as representative. However, for the 3 lead cable there can be differences in the hardware design that require it to be separately tested (see this MEDTEQ article on 3-Leads for more details).  

The test is heavily affected by noise. This is a result of the CMRR being degraded due to the high series impedance, or more specifically, the high imbalance in impedance. As this CMRR application note shows, CMRR is heavily dependent on the imbalance impedance. 

An imbalance of 620kΩ is 12 times larger than the CMRR test, so there is proportional degrading of the CMRR by the same factor of 12. This means for example that with a typical set up having 0.1mm (10µV@10mm/mV) mains noise for typical tests, would increase to 1.2mm of noise once the 620kΩ/4.7nF is in circuit. 

For the 0.67Hz test, the noise appears as a think line. It is possible consider the noise as an artefact and measure the middle point this think line (that is, ignore the noise). This is a valid approach especially as at 0.67Hz, there is usually no measurable reduction, so even increased measurement error from the noise, it is a clear "Pass" result. 

However, for the 40Hz test there is no line as such, and the noise is similar frequency resulting in beating, obscuring the result. And the result is often close to the limit. As such, the following is steps are recommended to minimise the noise:  

  • take extra care with the test environment, check grounding connections between the test circuit, the ECG under test, and the ground plate under the test set up
  • During measurement, touch the ground plate (this has often been very effective)
  • If noise still appears, use a ground plate above the test set up as well (experience indicates this works well)
  • Enable the mains frequency filter; this is best done after at least some effort is made to reduce the noise to using one or more of the methods above to avoid excessive reliance on the filter
  • Increase to a higher printing speed, e.g. 50mm/s 

Note that if the filter is used it should be on for the whole test. Since 40Hz is close to 50Hz, many simple filters have a measurable reduction at 40Hz. Since the test is proportional (relative), having the filter on does not affect the result as long as it is enabled for both the reference and actual measurement (i.e. with and without the test impedance). 

ECG Leads - an explanation

From a test engineer's point of view, it is easy to get confused with LEADS and LEAD ELECTRODES, because for a typical electrical engineer, "lead" and "electrode" are  basically the same thing. But there is more confusion here than just terminology. How do they get a "12 lead ECG" for a cable with only 10 leads? Why is it that many tests in IEC standards ask you to start with RA, yet the indication on the screen is upside down? Why is it that when you put a voltage on RA, a 1/3 indication appears on the V electrodes?

Starting with this matrix diagram, the following explanation tries to clear up the picture:

LEAD ELECTRODES are defined as the parts that you can make electrical connection to, such as RA, LA, LL, V1 and so on. On the other hand, LEADS are what the doctor views on the screen or print out.

There are a couple of reasons why these are different. Whenever you measure a voltage it is actually a measurement between two points. In a normal circuit there is a common ground, so we often ignore or assume this second reference point, but it's always there. Try and measure a voltage using one point of connection, and you won't get far.

ECGs don't have a common reference point, instead physicians like to see different "views" of the heart's electrical activity, each with it's own pair of reference points or functions of multiple points. One possibility would be to always identify the points of reference, but this would be cumbersome. Instead, ECGs use labels such as "Lead I" or "Lead II" to represent the functions.

For example "Lead I" means the voltage between LA and RA, or mathematically LA - RA. Thus, a test engineer that puts a positive 1mV pulse on RA relative to LA can expect to see an inverted (negative) pulse on Lead I.

Leads II and III are similarly LL-RA and LL-LA. 

The waveforms aVR, aVL and aVF are in effect the voltages at RA, LA and LL respectively, using the average of the other two as the second reference point.

Waveforms V1 ~ V6 (where provided) are the waveforms at chest electrodes V1 ~ V6 with the average of RA, LA and LL as the second reference point.

These 12 waveforms (Lead I, II, III, aVR, aVL, aVF, V1 ~ V6) form the basis of a "12 lead ECG".

Whether you are working with IEC 60601-2-27 or IEC 60601-2-51, you can refer to the diagram above or Table 110 in IEC 60601-2-51 which shows the relationship between LEAD ELECTRODES and LEADS.

Finally, you may ask what is RL (N) used for? The typical mistake is to assume that RL is a reference point or ground in the circuit, but this is not correct. In most systems, RL is not used for measurement. Rather it is used for noise cancellation, much like noise cancelling headphones, and is often call a "right leg drive". It senses the noise (usually mains hum) on RA/LA/LL, inverts and feeds back to RL. For testing IEC 60601-1, engineers should take note of the impedance in the right leg drive, as this tends to be the main factor which limits dc patient currents in single fault condition.

Exercise (test your understanding)

To check your understanding of the matrix, try the following exercise: if a 1mV, positive pulse (e.g. 100ms long) was fed to RA with all other inputs grounded, what would you expect to see on the screen for each lead? The answer is at the end of this page.

Other related information (of interest)

In years gone by, the relationship (matrix) above was implemented in analogue circuits, adding and subtracting the various input voltages. This meant that errors could be significant. Over time, the digital circuits have moved closer and closer to the inputs, and as well the accuracy of remaining analogue electronics has improved, which means it is rare to get any significant error in modern equipment. The newest and best equipment has a wide range high resolution input analogue to digital conversion very close to the input, allowing all processing (filtering as well as lead calculation) to be performed in software.

It is interesting to note that mathematically, even though there are 12 Leads, there are only 8 "raw" waveforms. Four of the 12 waveforms can be derived from the other 8, meaning they are just different ways of looking at the same information. For example, Lead III = Lead II - Lead I. It makes sense, since there are only nine points electrical connections used for measurement (remember, RL is not used for measurement), and the number of raw waveforms is one less than the number of measurement points (i.e. one waveform requires 2 measurement points, 2 waveforms requires at least 3 points, an so on). This is the reason why systems can use 8 channel ADC converters, and also why the waveform data used IEC 60601-2-51 tests (such as CAL and ANE waveforms) uses just 8 channels of raw data to create a full 12 Lead ECG.

Although the standard usually indicates that RA is the first lead electrode to be tested, if you want to get a normal looking waveform from a single channel source, it is best to put the output to LL (F) so that you get a positive indication on Lead II. Most systems default to a Lead II display, and often use Lead II to detect the heart rate. If your test system can select to put the output to more than one lead electrode, select LA and LL, which will give a positive indication on Lead I and Lead II (although Lead III will be zero).

Results of the exercise (answer)

If a +1mV pulse was applied to RA only, the following indications are expected on the screen (or printout). If you did not get these results or do not understand why these values occurred, go back and study the matrix relationships above. For the Lead electrodes, use RA = 1 and for all other use 0, and see what the result is.



Indication direction

Indication amplitude



















V1 ~ V6





CMRR Testing (IEC 60601-2-25, -2-27, -2-47)

Like EMC, CMRR testing is often considered somewhat of a black art in that the results are unpredictable and variable. This article attempts to clear up some of the issues by first looking at exactly how CMRR works in ECG applications and use of the RL drive to improve CMRR.

It also has a look at the importance of external noise, methods to eliminate and verify the set up is relatively free from external noise.

This application note is intended to support engineers that may already have some experience with CMRR testing but remained confused by variable results in individual set ups.

CMRR analysis from basics

CMRR is often considered a function of op-amp performance, but for the CMRR test in IEC/AAMI standards it turns out the indication on the ECG is mostly due to leakage currents passing through the 51k/47nF impedance.

First, let’s consider the basic test circuit:

For those wondering why the circuit shows 10V and 200pF rather than 20V and 100pF divider found in circuit found in IEC/AAMI standards, this arrangement is the “Thevenin equivalent” and can be considered identical. 

If this circuit was perfect, with the ECG inputs and gain element G floating with infinite input impedance, the 51k/47nF should have no effect and Lead I indication should be zero.

In practice, there will always be some small stray or deliberate capacitance in the system in the order 5 ~ 1000pF. This means the ECG inputs are not perfectly floating and small amounts of leakage will flow in the circuit.  

The main cause of this leakage is the capacitance between each input and shield or ground of the floating ECG circuit, and between that ECG shield/ground and the test system ground.

To understand how these influence the test it is best to re-arrange the circuit in a “long” fashion to appreciate the currents and current flow through the stray capacitance.

In this diagram, stray capacitance Ce-sg is added between the ECG electrode inputs and the ECG circuit ground (which is usually floating).

This capacitance is fairly high due to cable shielding and the internal electronics. Also each electrode has roughly the same stray capacitance. For example, a 12 lead diagnostic ECG measured around 600pF between RA and the shield, with a similar result for LA.

Capacitance Csg-tg between the ECG circuit ground (shield ground) and the test ground is also added.

This value can vary greatly, from as little as 5pF for a battery operated device with the cable well removed from the ground plane, to around 200pF for a mains operated device.

Lets assume Ce-sg are both 100pF, and Csg-tg is 10pF, and try to calculate the current that flows into the circuit. Although it looks complicated, it turns out the 51k/47nF is much smaller impedance compared to the stray capacitance, so as a first step we can ignore it. The total capacitance seen by the source is then a relatively simple parallel/series impedance calculation:  

                Ct = 1/(1/200+ 1/(100+100) + 1/10) = 9pF

We can see here that the largest impedance, in this case Csg-tg (shield to test ground), influences the result the most.


Next, we can calculate the total current flowing into the ECG:

                I = 10Vrms x 2π x 50Hz x 9pF = 28nArms

This seems very tiny, but keep in mind ECGs work of very small voltages.

The trick here is to realise that because Ce-sg is similar for RA and LA, this current will split roughly equally into both leads; around 14nA in our example.


RA has the imbalance of 51kΩ/47nF which has an impedance of Z = 40kΩ at 50Hz. When the 14nA flows thought this it creates 0.56mVrms between RA and LA. This is measured normally and on a 10mm/mV results in around 8mm peak to peak on Lead I of the ECG display.

To summarize, the 10Vrms will cause a small but significant amount of leakage to flow into the ECG circuit. This leakage will split roughly the same into each electrode. Any imbalance in the impedance of each electrode will cause a voltage drop which is sensed as a normal voltage and displayed on the ECG as usual.

In the above example, we can see that the capacitance Csg-tg between the ECG shield and the test ground had the largest effect on the result. We assumed 10pF, but increasing this to just 13pF would be enough to change this to a fail result. Many devices have 100pF or more; and the value can be highly variable due to the position of the shielded cable with respect to ground.

With such a small amount of highly variable capacitance having such a big effect, how can ECGs ensure compliance in practice?

The right leg drive

Most ECGs use a “right leg drive”, which is active noise cancellation and is similar to the methods used by noise cancellation headphones. Although noise “cancellation” implies a simple -1 feedback, it is often implemented a medium gain negative feedback loop, and sometimes with shield also driven at the +1 gain.

Regardless of the method, the basic effect is to absorb the leakage current through the RL electrode, which prevents it from creating a voltage across any impedance imbalance (51k/47nF).

In reality these circuits are not perfect, and in particular it is necessary to include a reasonable size resistor in the RL to prevent high dc currents going to the patient especially in fault condition. This resistor degrades the CMRR performance.

The residual indication on most ECGs (usually 3-7mm) is mostly a measure of the imperfection of the RL drive. This will be different for every manufacturer, but generally repeatable. Two test labs testing the same device should get similar results. Two samples of the same device type (e.g. production line testing) should give roughly the same results.

Since each RL drive system is different it can no longer be predicted how the system will react to changes in the position of the cable with respect to the ground plane. Test experience indicates that most ECGs with a RL drive, the indication reduces if the cable is closer to the test ground (Csg-tg capacitance is increased). With normal set ups, the variation is not big. In an extreme case, a test with 12 lead diagnostic ECG a portion of the cable was tightly wrapped in foil and the foil connected to the test ground. In this case the displayed signal to reduced by about 40%.

It is recommended that the ECG cable is loosely gathered and kept completely over the ground plane. Small changes in the cable position should not have a big effect and not enough to change a Pass/Fail decision. In case of reference tests the cable position might be defined in the test plan.

Systems without A Right leg drive

In general, all mains operated ECGs will employ a RL drive as the leakage will be otherwise too high.

In battery operated systems, some manufacturers may decide not use a RL drive.

Without a RL drive the analysis shows the test result will be directly proportional to the leakage current and hence highly sensitive to the cable position with respect to the test ground. The result will increase if the ECG device and cables are closer to test ground plane. This has been confirmed by experiment where a battery operated test sample without RL drive was shown to vary greatly with the sample and leads position with respect to ground plane, with both pass and fail results possible.

With the advent of wireless medical monitoring, there may be battery operated equipment intended for monitoring or diagnostic applications, together with inexperienced manufacturers that may not know the importance of the RL drive. Current standards (-2-25, -2-27) are not written well since they do not define what is done with the cable.

If a RL drive is not used, the above analysis indicates the intended use should be limited to being always worn on the patient and tested similar to IEC 60601-2-47. If the device has long cables and the recorder may be situated away from the patient, an RL drive should be used to avoid trouble.

For ambulatory equipment, the standard IEC 60601-2-47 specifies that the cable is wrapped in foil and connected to the common mode voltage, not the test ground. This is assumed to simulate the cable being close to the patient. This is expected to improve the result, as leakage will be much lower. The test voltage for ambulatory is also much smaller, at 2.8Vrms compared to 20Vrms. As such ambulatory equipment may pass without a RL drive.

External noise

In the actual CMRR test set up, the ECG electrodes are floating with around 10-15MΩ impedance to ground. This high impedance makes the circuit very susceptible to external noise, far more than normal ECG testing. The noise can interfere with the true CMRR result.  

Therefore for repeatable results, the test engineer must first set up to eliminate external noise as far as possible, and the test (verify) that there is no significant noise remaining.

To eliminate the noise the following steps should be taken:

  • Place the equipment under test (EUT), all cabling and the CMRR test equipment on an earthed metal bench or ground plane (recommended at least 1mm thick)
  • Connect the CMRR test equipment ground, EUT ground (if provided) and ground plane together and double check the connection using an ohm meter (should be <0.5Ω)
  • During the test, any people standing near the set up should touch the ground plane (this is an important step, as people make good aerials at 50/60Hz).

To check the set up has no significant noise:

  • Set up the equipment as normal, including the 20Vrms
  • Set RA lead with impedance (51k/47n), check normal CMRR indication appears (usually 3-8mm)
  • Turn the generator voltage off
  • Verify the indication on Lead I or Lead II is essentially a flat line at 10mm/mV. A small amount of noise is acceptable (e.g. 1mm) as long as the final result has some margin to the limit.

If noise is still apparent, a ground plane over the cables may also help reduce the noise. 

Typical Testing Results

Most indications for the 20V tests are in the range of 3-7mm. An indication that is lower or higher than this range may indicate there problem with the set up.

Indications are usually different for each lead which is expected due to the differences in the cable and trace layout in the test equipment, test set up and inside the equipment under test. Therefore, it is important to test all leads. 

The 300mVdc offset usually has no effect on the result. However, the equipment has to be properly designed to achieve this result - enough head room in the internal amplifiers. So it is again important to perform the test at least for representative number of leads.

If the test environment is noisy, there may be "beating" between the test signal frequency (which is usually pretty accurate) and real mains frequency, which is not so accurate. This can be eliminated by taking special precautions with grounding and shielding for the test area. Solid metal benches (with the bench connected to the test system ground) often make the best set up. 

And that 120dB CMRR claim? 

Some ECG manufacturers will claim up to 120dB CMRR, a specification which is dubious based on experience with real ECG systems. The requirement in standards that use the 10V test is effectively a limit of 89dB  (= 20 log (0.001 / (2√2 x 10)). A typical result is around 95dB. Although it might not seem much between 95dB and 120dB, in real numbers it is a factor of about 20. 

It is likely that the claim is made with no imbalance impedance - as the analysis above shows, the imbalance creates the common mode indication, and without this imbalance most floating measurement systems will have no problem to provide high CMRR. Even so, in real numbers 120dB is a ratio of a million to 1, which makes it rather hard to measure. So the claim is at best misleading (due to the lack of any imbalance) and dubious, due to the lack of measurement resolution. Another challenge for standards writers?     

ECG Filters

ECG filters can have a substantial effect on the test results in IEC 60601-2-25, IEC 60601-2-27 and IEC 60601-2-47. In some clauses the standard indicates which filter(s) to use, but in most cases, the filter setting is not specified. One option is to test all filters, but this can be time consuming. Also, it is not unusual to find that some tests fail with specific filter settings. This section is intended to give some background on the filters and the effect of filters, so test engineers can decide which filter settings are appropriate.

Most test engineers covered filters at some point in their education, but that knowledge may have become rusty over time, so the following includes some information to brush up on filter theory while heading into the specifics of ECG filters.

Section 1: The technology behind filters

What is a filter?

In general, filters try to remove unwanted noise. Especially in ECG work, the signal levels are very small (around 1mV), so it is necessary to use filtering to remove a wide range of noise. This noise may come from an unstable dc offset from electrode/body interface, muscle noise, mains hum (50/60Hz), electrical noise from equipment in the environment and from within the ECG equipment itself, such as from internal dc/dc converters.

A filter works by removing or reducing frequencies where noise occurs, while allowing the signal frequency through. This can be done in either hardware or software. In modern systems, the main purpose of hardware filtering is to avoid exceeding the limits of the analogue system, such as opamp saturation and ADC ranges. Normally a 1mV signal would be amplified around 100-1000 times prior to ADC sampling, if this signal had even 10mV of noise prior to amplification, we can expect amplifiers to saturate. The main limitation of hardware filters is that they rely on capacitors, the value of which cannot be controlled well both in production and in normal use. Thus software filtering is usually relied on for filter cut-off points that can be controlled accurately, allowing also advanced filter models and user selected filters to be implemented. 

What are typical types of ECG filtering? Why are there different filters?

Ideally, a filter should remove noise without affecting the signal we are interested in. Unfortunately, this is rarely possible. One reason is that the signal and noise may share the same frequencies. Mains noise (50/60Hz), muscle noise and drift in dc offsets due to patient movement all fall in the same frequency range as a typical ECG. Another problem is that practical filters normally don't have a sharp edge between the "pass" band and the "cut" band. Rather there is usually a slow transition in the filters response, so if the wanted and unwanted signals are close we may not be able to remove the noise without removing some of the desired signal.

The result is that filters inevitably distort the signal frequency. The image right shows the distortion of the ANE20002 waveform from IEC 60601-2-25 with a typical "monitor" filter from 0.67Hz to 40Hz. A balance has to be found between removing noise and preserving the original signal. For different purposes (monitoring, intensive care, diagnostic, ambulatory, ST segment monitoring etc) the balance shifts, so we end up with a range of filters adjusted to get the best balance. Some common examples of ECG filters are:

Diagnostic:   0.05Hz ~ 150Hz    
Widest for diagnostic information, assumes a motionless, low noise environment

Ambulatory, patient monitoring:    0.67Hz ~ 40Hz 
Mild filtering for noisy environment, principally to detect the heart rate

ST segment:  0.05Hz ~    
Special extended low frequency response for ST segment monitoring (more detail below)

Muscle, ESU noise:   ~ 15Hz   
Reduced higher frequency response to eliminate muscle noise and other interference such as ESUs

While ECGs could be referred to as using a band pass filter, the upper and lower frequencies of the pass band are sufficiently apart that we can discuss them seperately as low pass and high pass filters.

What is a low pass filter? What distortion is caused by low pass filtering?

A low pass filter is often found in electronic circuits, and works by reducing high frequency components. The most common form of a hardware low pass filter is a simple series resistor / capacitor: at low frequencies the capacitor is high impedance relative to the resistor, but as the frequency increases the capacitor impedance drops and output falls. A circuit with only one resistor/capacitor is a "single pole filter". Due to origins in audio work and similar fields, filters are normally specified by the frequency at which there is a "3dB reduction", or where the output voltage is around 71% (0.707) of the input. While this may sound large, in the audio field the dynamic range is so large that a log scales are required, and on this scale 3dB reduction (30%) is not so big. For a large dynamic range, units of decibels (dB) are more convenient. Decibels originated in power, using simple scale of 10 log10(Pout / Pin). In electronics, measurement of voltage is more common, thus we end up with 20 log10(Vout / Vin). The factor of 20 rather than 10 reflects the square relationship between voltage and power, which in the log world is an additional factor of 2.    

The use of log scales can be misleading. Graphically in the log/log scale, the output of a single pole filter is basically 1:1 (100%) in the "pass band", and then drops of steeply as the frequency increases, quickly reaching levels of 1% (0.01) and lower.  

However, if we look at a graph using a normal scale (non-log), we see that around the frequency of interest, the cut of is actually pretty slow. For example, for a 40Hz filter, at 20Hz there will still be more than 10% reduction, and at 100Hz, still 37% of the signal is getting through. When testing an ECG's filter response and other characteristics, is it common to see effects due to filters above and below the cut off frequencies.

In software, filters can be used which closely approximate hardware filters, but other complex forms are possible. Sharper cut off between the pass band and cut band can also be achieved. Great care is needed with software filters as unexpected results can easily occur due to the interplay between sampling rates and the chosen methodology.  

The distortion caused by a hardware (or equivalent software) single pole low pass filter is easy to visualise: it essentially dampens and slows the waveform, much like suspension in a car. The following graph shows the effect of a 40Hz monitoring filter on 100ms rectangle and triangle pulses. For the triangle it is interesting to note that there is about a 5% reduction in the peak measured, and also a small delay of around 3ms.

What is a high pass filter? What are the effects?

High pass filters are obviously the opposite of a low pass filters. In hardware, a single pole filter can be made out of a capacitor in series with a resistor. The corner frequency is the same, and the frequency response is a mirror image (vertical flip) of the low pass filter.

The terminology associated with ECG high pass filters can be confusing: while the filter is correctly termed a "high pass filter", it affects the low frequency response, around the 0.05Hz to 1Hz region. So it is easy to get mixed up between "high" and "low".

The main intention of a high pass filter in ECG work is to remove the dc offset which in turn is largely caused by the electrode/gel/body interface. Unstable voltages of up to 300mVdc can be produced. In diagnostic work, the patient can be asked to stay still so as to reduce these effects, allowing the filter corner to be reduced down to 0.05Hz. For monitoring and ambulatory use, a 0.67Hz corner is common.

For long term periodic waveforms the main effect is to shift or keep the waveform around the centerline, known as the "baseline" in ECG. This is the same as using the AC mode on an oscilloscope to view only ac noise of 50mVpp on a 5Vdc supply rail. Most test engineers have little problem to understand this side of high pass filters.  

However, for short term pulses, the effects of high pass filters on waveforms are not so easy to visualise. In particular, it is possible to get negative voltages out of a positive pulse waveform, and also peak to peak values exceeding the input. These effects cannot occur with a low pass filter. The hardware filter circuit shown just above, together with the graph below can help to understand why this happens. Initially the capacitor has no charge, so that when a step change (1V) is applied, the full step is transferred to the output. Then the capacitor slowly charges according to the circuit's time constant. For a filter with 0.67Hz, after 100ms, the capcitor is charged to around 0.34V. When the input suddenly drops to 0V, the capacitor remains charged at 0.34V, but the polarity is negative with respect to Vout. The output voltage is Vout = Vin - Vc = 0 - 0.34 = -0.34V. As long as the input remains at 0V, the capcitor then slowly discharges back towards 0V. In this way we can get negative voltages from a positive pulse, a peak to peak voltage of 1.34V (exceeding the input), and finally long slow time constants resulting from short impulses.  

This long time constant can cause problems in viewing the ECG trace after large overloads, such as during defibrillator pulses or a temporary disconnected lead. A 0.67Hz high pass filter has a 0.25s time constant, which although is short can still take time since the overloads are in the 1V level, 1000 times higher than normal signals. For these reasons, ECGs are often provided with "baseline reset" or "de-blocking" function to reset the high pass filter. Typically this is an automated function which in hardware filtercan be done by shorting the capacitor (e.g. analogue or FET switch), or in software filters is simply clearing a result back to zero. 

Diagnostic filters and other filters that go down to 0.05Hz have a much slower time constant, so it can take 10-15s for the signal to become visible again. Even after an automated baseline reset there may be residual offsets of 5-50mV which keep the signal off the screen. This can be a serious risk if such filters are used in intensive care patient monitoring. Patient monitors are often provided with both diagnostic and monitoring filters, and while they pass defibrillator and 1V 50/60Hz overload tests with the monitoring filter, they fail when tested with a diagnostic filter setting. This is a subject which can cause conflict as the standard does not define which filter to use, and manufacturers often argue that only the monitor filter should be tested. However, basic risk management indicates that regardless of the filter setting, the baseline reset should work effectively. It is fairly obvious that such a filters with 0.05Hz would not be selected for defibrillation, however, it is also unlikely that if the patient monitor was already set to diagnostic mode prior to an emergency situation , we cannot reasonably expect the operator to remember or have the time to mess around changing filter settings. Also, the technology to detect and reset the baseline after overloads is well established.

ST filters are also common in patient monitoring and create a similar problem. The purpose of the filter is to preserve the "ST segment" which occurs between the QRS pulse and T wave and can be an important diagnostic indicator. The following graph shows how normal monitoring ECG high pass filter of 0.67Hz on the CAL20160 waveform from IEC 60601-2-25 (+0.2mV elevated ST segment) essentially removes the ST segment:

If we reduce the low frequency response (high pass filter) down to 0.05Hz, we can see that the ST segment remains largely undistorted, allowing diagnostic information to be retained:

Notch filters (mains hum filters, AC filter, 50/60Hz)

Notch filters combine both high and low pass filters to create a small region of frequencies to be removed. For ECGs, the main target is to remove 50Hz or 60Hz noise. Because mains noise falls in the region of interest (especially for diagnostic ECGs), the setting of "AC filter" is usually optional. ECG equipment already contains some ability to reject mains noise even without a filter (see right leg drive) so depending on the amount of AC noise in the environment, an AC filter may not be required. A good check of your ECG testing location is to compare the signals with and without the AC filter on.

Some systems automatically detect the mains frequency, others are set by the user or service personnel, while others use a single notch filter covering both 50/60Hz.

High "quality" notch filters can be created in software that target only 50 or 60Hz, but the drawback of these filters is they can create unusual ringing especially to waveforms with high rates of change. IEC 60601-2-51 has a special waveform (ANE20000) which confirms that the extent of ringing is within reasonable limits.

Similar to the diagnostic filter, the question again arises as to whether patient monitors should pass tests with or without the AC filter. In particular this causes problems with the 40Hz high frequency response requirement, as some systems may fail this response with a 50Hz AC filter on. There is no simple answer for this: 40Hz and 50Hz are very close, so to comply with the 40Hz requirement with a 50Hz notch filter implies advanced multipole filtering. But multipole filters have risks of distortion such as ringing. On the other hand, use of AC filters can be considered "normal condition", so to argue that a test is not required with the AC filter on implies that the 40Hz frequency response is not really important, which would raise the question what upper frequency response is important. ANSI/AAMI (US) standards have an upper limit of 30Hz for patient monitors, which also complicates the situation.

Ultimately, the final decision would require a careful study of the effects of the AC filters on waveforms found in real clinical situations, which also depends in the intended purpose. In particular neonatal waveforms are likely to have higher frequency components, so the high frequency response including AC filters will have the greatest impact only if neonatal patients are included in the intended purpose. The following images show the effects of 40Hz and 30Hz single pole filters on IEC 60601-2-51 waveform CAL20502 (intended to simulate neonatal ECGs). As the images show, the effects are not insignificant. Both filters reduce the peak to peak indication, with the 30Hz filter around 20%, which may be exceeding reasonable limits. However, of course these are single pole filter simulations, which would not relfect the response of more complex filter systems.  

Notes on advanced filtering

The simulations above are based on simple single pole filters, which distort the signal in predictable ways and are easy to simulate. Complex multipole and digital filters can have far better responses but there are risks of substantial overshoots and ringing. Experience from testing indicates that manufacturers tend to prefer simple filters, but occasionally use more complex filters where strange results in testing are possible. These results may or may not representative of the real world because the test signals often contain frequencies that don't exist in the real world, such as small digital steps caused by arbitrary waveform generators, or rectangle pulses with excessively fast rise times. This needs to be kept in mind during testing and discussed with the manufacturer.

Section 2: Particular requirements from standards affected by filters

Sensitivity, accuracy of leads, accuracy of screen and printing, similar tests

For tests involving sensitivity (e.g. confirming 10mm/mV within ±5%) and accuracy of lead calculations (such as Lead I = RA - LA), it makes sense to use diagnostic filter with the AC filter off. The nature of these tests is such that filters should not impact the result, with the effects of filters being handled seperately. The widest frequency response ensures that the waveforms viewed on the screen are essentially the same as the input waveforms, avoiding some complications due to waveform distortion which are irrelevant to the tests. This assumes that the test environment is sufficiently "quiet" so that mains and other noise does not also influence the result.

Common Mode Rejection Ratio

As IEC standards point out, the CMRR test should be performed with the AC filter off, if necessary by special software. If avaliable, a patient monitor should be tested using the widest (diagnostic) filter mode, which is worst case compared to monitor mode. One point to note is that ANSI/AAMI standards (at least, earlier editions) do not require the AC filter to be off, a key difference to the tests in IEC standards.

Input impedance test

Due to the high imbalance in one lead (620k/4.7nF), the input impedance test is particularily susceptable to mains noise. Since this is a ratiometric test, the filter setting should not affect the result. If possible, the user should select the mains notch filter to be on, and use the monitoring mode. Other filter settings (like muscle, ESU) might reduce the noise further, but they may also make it difficult to measure at 40Hz  as the signal will be substantially attenuated. 

Frequency response test

For frequency response tests, including the 200ms/20ms triangle impulse test, obviously all filters should be tested individually. However, there may be discussions as indicated above as to whether compliance is necessary for all settings, which in turn may be based on clinical discusssion. For example, it is obvious that special filters in highly noisy environments (e.g. muscle, ESU) may not meet the 40Hz high frequency response requirment from IEC 60601-2-27. Test labs should simply report the results. For regulatory purposes, manufacturers should discuss the clinical impact where appropraite. For example, a muscle filter with a cut off of 15Hz seems clearly inappropriate for use with neonates. 

For IEC 60601-2-27 (0.67Hz to 40Hz), practical tests found that some manufacturers follow the normal practice in frequency response testing of using the input as the reference. For example setting the input to exactly 1mVpp (10mm) and then measuring the output. While this is logical, the standard requires that the output at 5Hz is used as the reference point. In some cases, the 5Hz output can be significantly higher than the input as the result of multipole filters, leading to differences between manufacturer test results in independent laboratory test results.

For IEC 60601-2-25, frequency sweeps up to 500Hz using digital based systems usually finds some point where beating occurs, as a result of the sample rate of the digital function genorator being is a multiple or near multiple of the ECG's sampling rate. For this reason, it is always useful to have a back up analogue style function genorator on hand to verify the frequency response.

Pacemaker indication

Most modern ECGs use a blanking approach to pacemaker pulses: automatic detection of the fast edge of the pacing spike, ignoring the data around the pulse and then replacing the pulse with an artificial indication on the ECG screen or record. If this approach is taken, the filter settings usually do not affect the test results. However, some systems allow the pulse through to the display. In this case, the filter settings can dramatically affect the result. Consult the operation manual prior to the test to see if any special settings are necessary for compliance.  

Low frequency impulse response test (3mV 100ms)

The low frequency impulse response test is only intended where the frequency response extends down to 0.05Hz. For patient monitors and ambulatory ECGs, this will typically only apply for special settings such as diagnostic filters or ST-segment analysis. There appears to be an error in IEC 60601-2-47 since it requires the test for all modes, but it is obvious that filters that do not extend down to 0.05Hz cannot pass the test.

Simulations with a single pole filter 0.05Hz have found that the results just pass the tests in IEC 60601-2-27 and IEC 60601-2-47, with an overshoot of 93uV and a slope of 291uV/s, compared to the limits of 100uV and 300uV/s in the standards. It appears that IEC 60601-2-51 cannot be met with a single pole filter, as it has a slope requirement of 250uV/s. The rationale in the standard indicates that this is intentional. It is very difficult if not impossible to confirm compliance based on inspection of print outs as the values are very small, so it may require digital simulations and assistance from the manufacturer, with the full system test (analogue signal through to the printout) being used only for confirmation. The following graphs show simulated responses for 0.05Hz single pole filter, both overall and a close up of the overshoot

 Any questions or comments, please feel free to contact




IEC 60601-2-25 Clause - Goldberger and Wilson LEADS

In a major change from the previous edition (IEC 60601-2-51:2003), this standard tests the Goldberger and Wilson LEAD network using CAL waveforms only. There are some concerns with the standard which are outlined below: 

  • the standard does not indicate if the tests must be performed by analogue means, or if optionally digital tests are allowed as indicated in other parts of the standard. It makes sense to apply the test in analogue, as there is no other test in the standard which verifies the basic accuracy of sensitivity for the complete system (analogue and digital).
  • The CAL (and ANE) signals are designed in a way that RA is the reference ground (in the simulation data, RA is always zero; in the circuit recommended in IEC 60601-2-51, RA is actually connected to ground). This means that an error on RA cannot be detected by CAL or ANE signals. The previous standard was careful to test all leads individually, including cases where a signal is provided only to RA (other leads are grounded), ensuring errors on any individual lead would be detected. 
  • The allowable limit is 10%. This is a relaxation from IEC 60601-2-51, conflicts with the requirement statement in Clause and also requirements for voltage measurements in Clause, all of which use 5%. Furthermore, many EMC tests refer to using CAL waveforms with the criteria from Clause (5%), not the 10% which comes from this clause.  

A limit of 5% makes sense for diagnostic ECGs and is not difficult with modern electronics and historically has not been an issue. There is no explanation where the 10% comes; at a guess the writers may have trying to separate basic measurement sensitivity (5%) from the network accuracy (5%). In practice, it makes little sense to separate these out as ECGs don't provide access to the raw data from each lead electrode, only the final result which includes both the sensitivity and the network. As such we can only evaluate the complete system based on inputs (lead electrodes, LA, LL, RA etc) and outputs (displayed LEAD I, II, III etc).  

As mentioned above, there is no other test in IEC 60601-2-25 which verifies the basic sensitivity of the ECG. Although sensitivity errors may become apparent in other tests, it makes sense to establish this first as a system, including the weighting network, before proceeding with other tests. While modern ECGs, from quality manufacturers and designed specifically for for diagnostic work generally have little problem for 5%, experience indicates that lower quality manufacturers and in particular multipurpose devices (e.g. patient monitor with diagnostic functions) can struggle to meet basic accuracy requirement for sensitivity. 

IEC 60601-2-25 Clause Indication of Inoperable ECG

This test is important but has a number of problems in implementation. To understand the issue and solution clearly, the situation is discussed in three stages - the ECG design aspect the standard is trying to confirm; the problems with the test in the standard; and finally a proposed solution. 

The ECG design issue

Virtually all ECGs will apply some opamp gain prior to the high pass filter which removes the dc offset. This gain stage has the possibility to saturate with high dc levels. The point of saturation varies greatly with each manufacturer, but is usually in the range of 350 - 1000mV. At the patient side a high dc offset is usually caused by poor contact at the electrode site, ranging from an electrode that is completely disconnected through to other issues such as an old gel electrode. 

Most ECGs detect when the signal is close to saturation and trigger a "Leads off" or "Check electrodes" message to the operator. Individual detection needs to be applied to each lead electrode, and both positive and negative voltages, this means that there are up to 18 different points (LA, RA, LL, V1 - V6). Due to component tolerances, the points of detection in each lead often vary by around 20mV  (e.g. LA points might be +635mV, -620mV, V3 might be +631mV, -617mV etc). 

If the signal is completely saturated it will appear as a flat-line on the ECG display. However, there is a small region where the signal is visible, but distorted (see Figure 1). Good design ensures the detection occurs prior to any saturation. Many ECGs automatically show a flat line once the "Leads Off" message is indicated, to avoid displaying a distorted signal. 

Problems with the standard

The first problem is the use of a large ±5V offset. This is a conflict with the standard as Clause states that ECGs only need to withstand up to ±0.5V without damage. Modern ECGs use ±3V or less for the internal amplifiers, and applying ±5V could unnecessarily damage the ECG. 

This concern also applies to the test equipment (Figure 201.106). If care is not taken, the 5V can easily damage the precision 0.1% resistors in the output divider and internal DC offset components.  

Next, the standard specifies that the voltage is applied in 1V steps. This means it is possible to pass the test even though equipment fails the requirement. For example an ECG may start to distort at +500mV, flatline by +550mV, but the designer accidentally sets the "Leads Off" signal at +600mV. In the region of 500-550mV this design can display a distorted signal without any indication, and from 550-600mV is confusing to the operator why a flat line appears. If tested with 1V steps these problem regions would not be detected and a Pass result would be recorded. 

Finally the standard allows distortion up to 50% (a 1mV signal compressed to 0.5mV). This is a huge amount of distortion and there no technical justification to allow this given the technology is simple to ensure a "Leads Off" message appears well before any distortion. The standard should simply keep the same limits for normal sensitivity (±5%).


In practice, it is recommended that a test engineer start at 300mV offset and search for the point where the message appears, reduce the offset until the message is cleared, and then slowly increase again up to the point of message while confirming that no visible distortion occurs (see Figure 2). The test should be performed in both positive and negative directions, and on each lead electrode (RA, LA, LL, V1 to V6). The dc offset function in the Whaleteq SECG makes this test easy to perform (test range up to ±1000mV), but the test is also simple enough that an ad-hoc set up is easily prepared.  

Due to the high number of tests, it might be temping to skip some leads on the basis that some are representative. Unfortunately, experience indicates that manufacturers sometimes deliberately or accidentally miss some points on some leads, or set the operating point to the wrong level, such that distortion is visible prior to the message appearing. As such it is recommended that all lead electrodes are checked. Test engineers can opt for a simple OK/NG record, with the operating points on at least one lead kept for reference. Detailed data on the other leads might be kept only if they are significantly different. For example, some ECGs have very different trigger points for chest leads (V1 - V6). 

Due to the nature of electronics, any detectable distortion prior to the "Leads Off" message should be treated with concern, since the point of op-amp saturation is variable. For example one ECG may have 10% distortion at +630mV while sample might have 35% distortion. Since some limit should apply (it is impossible to detect "no distortion") It is recommended to use a limit of ±5% relative to a reference measurement taken with no dc offset. 

The right leg

The right leg is excluded from the above discussion: in reality the right leg is the source of dc offset voltage - it provides feedback and attempts to cancel both ac and dc offsets; an open lead or poor electrode causes this feeback to drift towards the internal rail voltage (typically 3V in modern systems). This feedback is via a large resistor (typically 1MΩ) so there is no risk of damage (hint to standards committees - if 5V is really required, it should be via a large resistor).

More research is needed on the possibility, effects and test methods for RL. It is likely that high dc offsets impact CMRR, since if the RL drive is pushed close to rail voltage, it will feedback a distorted signal preventing proper noise cancellation. At this time, Figure 201.106 is not set up to allow investigation of an offset to RL while providing signals to other electrodes which is necessary to detect distortion. For now, the recommendation is to test RL to confirm that at least an indication is provided, without confirming distortion.

Figure 1: With dc offset, the signal is at first completely unaffected, before a region of progressive distortion is reached finally ending in flat line on the ECG display. Good design ensures the indication to the operator (e.g. "LEADS OFF") appears well before any distortion

Figure 2: The large steps in the standard fail to verify that the indication to the operator appears before any distortion. Values up to 5V can also be destructive for the ECG under test and test equipment. 

Figure 3; Recommended test method; search for the point when the message is displayed, reduce until message disappears, slowly increase again check no distortion up to the message indication. Repeat for each lead electrode and both + and - direction.

IEC 60601-2-25 Clause Defibrillator Protection

General information on defibrillator testing can be found in this 2009 article copied from the original MEDTEQ website.

One of the significant changes triggered by the IEC 60601-2-25 2011 edition is the inclusion of the defibrillator proof energy reduction test via the general standard (for patient monitors, this test already existed via IEC 60601-2-49). Previously, diagnostic ECGs tended to use fairly low impedance contact to the patient, which helps to improve performance aspects such as noise. The impact of the change is that all ECGs will require higher resistors in series with each lead, as detailed in the above article. The higher resistors should trigger retesting for general performance, at least for a spot check.

Experience from real tests has found that with the normal diagnostic filter (0.05Hz to 150Hz), the baseline can take over 10s to return, exceeding the limit in the standard. Although most systems have automated baseline reset (in effect, shorting the capacitor in an analogue high pass filter, or the digital equivalent), the transients that occur after the main defibrillator pulse can make this difficult for the system to know when the baseline is sufficiently stable to perform a reset.  The high voltage capacitor used for the main defibrillator pulse is likely to have a memory effect causing significant and unpredictable baseline drift well after the main pulse. If a reset occurs during this time, the baseline can easily drift off the screen, and due to the long time constant of the 0.05Hz filter, can take 15-20s to recover. 

The usual workaround is that most manufacturers declare in the operation manual that during defibrillation special filters should be used (e.g. 0.67Hz). The issue raises the question of why diagnostic ECGs need to have defibrillator protection, and if so, how this is handled in practice. If defibrillator protection is really necessary, sensible solutions may involve the system automatically detecting a major overload and switching to a different filter for a short period (e.g. for 30s). It is after all an emergency situation: expecting the operator to have read, understand and remember a line from the operation manual, and as well have the time and presence of mind to work through touch screen menu system to enable a different filter setting while at the same time performing defibrillation on the patient is a little bit of a stretch. 

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IEC 60601-2-25 CSE database - test experience

The standard IEC 60601-2-25:2011 includes tests to verify the accuracy of interval and duration measurements, such as QRS duration or the P-Q interval.

These tests are separated into the artificial waveforms (the CTS database) and the biological waveforms (the CSE database). The CSE database is available on CD-ROM and must be purchased from the INSERM (price $1500, contact Paul Rubel,[1].

In principle, the database tests required by IEC 60601-2-25:2011 should be simple: play the waveform (in digital or analogue form), compare the data from the equipment under test to the reference data. In practice, there are some considerations and complications. This document covers some of the issues associated with the CSE database.  

First, it should be confirmed that the equipment under test actually measures and displays global intervals, rather than intervals for a specific lead. As stated in Annex FF.2 of the standard:

“The global duration of P, QRS and T are physiologically defined by the earliest onset in one LEAD and the latest offset in any other LEAD (wave onset and offset do not necessarily appear at the same time in all LEADS because the activation wavefronts propagate differently).”

Global intervals can be significantly different to lead intervals. This becomes evident from the first record of the database (#001), where the reference for the QRS duration is 127ms, while the QRS on LEAD I is visibly around 100ms. The following image is from the original “CSE Multilead ATLAS” analysis for recording #001 shows why: the QRS onset for Lead III (identified with the line and sample number 139) starts much earlier than for Lead I.

If the equipment under test does not display global intervals, it is not required to test using the CSE database to comply with IEC 60601-2-25.

The next aspect to be considered is whether to use waveforms from the MO1 or MA1 series.

The MO1 series is the original recording, and contains 10s with multiple heart beats. Each heart beat is slightly different, and the reference values are taken only for a specific heart beat (generally the 3rd or 4th beat in the recording). The beat used for analysis can be found from the sample number in the file “MRESULT.xls” on the CD-ROM[2]. The MO1 recordings are intended for manufacturers using the software (digital) route for testing their measurement algorithms. Many ECGs perform the analysis by averaging the results from several beats, raising a potential conflict with the standard since the reference values are for a single beat only. It is likely that the beat to beat variations are small and statistically insignificant in the overall analysis, as the limits in the standard are generous. However manufacturers investigating differences in their results and the reference values may want to check other beats in the recording.

The MO1 files can played in analogue form but there are two disadvantages: one is to align the equipment under test with the reference beat, the second is looping discontinuities. For example Record 001 stops in the middle of a T-wave, and Lead V1 has a large baseline drift. If the files are looped there will be a large transient and the potential for two beats to appear together; the ECG under test will have trouble to clear the transient events while attempting to analyze the waveforms. If the files are not looped, the ECG under test may still have trouble: many devices take around 5s to adjust to a new recording, by which time the reference beat has already passed.

The MA1 series overcomes these problems by isolating the selected beat in the MO1 recording, slightly modifying the end to avoid transients, and then stitching the beats together to make a continuous recording of 10 beats. The following image superimposes the MA1 series (red) on top of the MO1 series (purple) for the 4th beat on Lead I. The images are identical except for a slight adjustment at the end of the beat to avoid the transient between beats:

 The MA1 series is suitable for analogue and digital analysis. Unlike MO1 files which are fixed at 10.0s, the MA1 files contain 10 whole heart beats, so the length of the file varies depending on the heart rate. For example, record # 001 has a heart rate around 63bpm, so the file is 9.5s long. Record 053 is faster at 99bpm, so the file is only 6s long. As the file contains whole heart beats, the file can be looped to allow continuous play without limit. There is no need to synchronize the ECG under test, since every beat is the same and the beat is always the reference beat.

The only drawback of the MA1 series is the effect of noise, clearly visible in the above original recording. In a real recording, the noise would be different for each beat, and helps to cancel out errors if averaging is used. For manual analysis (human), the noise is less of a concern as we can visually inspect all leads simultaneously, from this we can generally figure out the global onset even in the presence of noise. Software usually looks at each lead individually and can be easily tricked by the noise. This is one reason why ECGs often average over several beats. Such averaging may not be effective for the MA1 series since the noise on every beat is the same.

Finally, it should be noted that the CSE database contains a large volume of information much of which is irrelevant for testing to IEC 60601-2-25. Sorting through this information can be difficult. Some manufacturers and test labs, for example, have been confused by the file “MRESULTS.xls” and attempted to use the reference data directly. 

In actual case, file “MRESULTS.xls” does not contain the final reference values used in IEC tests. They can be calculated from the raw values (by selective averaging), but to avoid errors it is best to use the official data directly.

Most recent versions of the CD-ROM should contain a summary of the official reference values in three files (all files have the same data, just difference file format):

  • IEC Biological ECGs reference values.pdf
  • IEC Biological ECGs reference values.doc
  • CSE_Multilead_Library_Interval_Measurements_Reference_Values.xls

If these files were not provided in the CD-ROM, contact Paul Rubel (

[1] The CSE database MA1 series are embedded in the MEDTEQ/Whaleteq MECG software, and can be played directly without purchasing the CD-ROM. However the CD-ROM is required to access the official reference values.

[2] For record #001, the sample range in MRECORD.xls covers two beats (3rd and 4th). The correct beat is the 4th beat, as shown in the original ATLAS records, and corresponds to the selected beat for MA1 use.

IEC 60601-2-25 Overview of CTS, CSE databases

All ECG databases have two discrete aspects: the digital waveform data, and the reference data. The waveform data is presented to the ECG under test, in either analogue or digital form (as allowed by the standard), and the ECG under test interprets the waveform data to create measured data. This measured data is then compared against the reference data to judge how well the ECG performs. These two aspects (waveform data, reference data) need to be considered separately. This article covers the databases used in IEC 60601-2-25.  

CTS Database

The CTS database consists of artificial waveforms used to test for automated amplitude and interval measurements. It is important to note that the standard only applies to measurements that the ECG makes: if no measurements are made, no requirement applies; if only the amplitude of the S wave in V2 is measured, or duration of the QRS of Lead II, that is all that needs to be tested. In the 2011 edition the CTS database is also used for selected performance tests, some of which needs to be applied in analogue form. 

All the CAL waveforms are identical for Lead I, Lead II, V1 to V6, with Lead III a flatline, aVR inverted and aVL, aVF both half amplitude, as can be predicted from the ECG Leads relationship. The ANE waveforms are more realistic, with all leads having similar but different waveforms. A key point to note with the ANE20000 waveform is the large S amplitude in V2, which usually triggers high ringing in high order mains frequency filters - more on that on another page.  

The CTS waveform data is somewhat of a mystery. in 2009 MEDTEQ successfully bought the waveforms from Biosigna for €400, but that organisation is now defunct (the current entity bears no relation). The standard states that the waveforms are part of the CSE database and available from INSERM, but this information is incorrect, INSERM is responsible for CSE only. According to Paul Rubel (INSERM), the CTS database was bought by Corscience, but their website contains no reference to the CTS database, nor where or how to buy it.  

Adding to the mystery, In the 2005 edition of IEC 60601-2-25 the CTS reference data was mentioned in the normative text but completely missing in the actual appendixes. The 2011 edition finally added the data but there are notable errors. Most of the errors are easily detected since they don't follow the lead definitions (for example, data for Lead I, II and III is provided, and this must follow the relation Lead III = Lead II - Lead I, but some of the data does not),  

Almost certainly, the situation is affected by a moderately wide group of individuals associated with the big manufacturers that are "in the know" and informally share both the waveform and reference data with others that are also "in the know" - otherwise it seems odd that the errors and omissions would persist. Those of us outside the the group are left guessing. The situation is probably illegal in some countries - standards and regulations are public property and the ability to verify compliance should not involve secret knocks and winks. 

The good news is that thanks to Biosigna, MEDTEQ and now Whaleteq has the CTS database embedded in the MECG software. And the reference data is now in the standard. This provides at least one path for determining compliance. We are not permitted to release the digital data. 

The experience from actual amplitude tests has been good. Most modern ECGs (software and hardware) are fairly good at picking out the amplitudes of the input waveform and reporting these accurately and with high repeatability. Errors can be quickly determined to be either:

  • mistakes in the reference data (which are generally obvious on inspection, and can be double checked against the displayed waveforms in MECG software);
  • due to differences in definitions between the ECG under test and those used in the standard;
  • due to the unrealistic nature of the test waveforms (for example, CAL50000 with a QRS of 10mVpp still retains a P wave of just 150µV); or
  • an actual error in the ECG under test  

For CTS interval measurements, results are mixed. Intervals are much more difficult for the software as you need to define what is a corner or edge (by comparison, peak is peak, it does not need a separate definition). Add a touch of noise, and the whole interval measurement gets messy. Which is probably why the standard uses statistical analysis (mean, deviation) rather than focusing on any individual measurement. Due to the statistical basis, the recommendation here is to do the full analysis first before worrying about any individual results. 

CSE Database

For the CTS database, the standard is actually correct to refer to INSERM to obtain both the waveform and reference data. The best contact is Paul Rubel ( Unlike CTS, the CSE database uses waveforms from real people and real doctors were involved in measuring the reference data. As such it is reasonable to pay the US$1500 which INSERM requests for both the waveforms and reference data,

The MECG software already includes the CSE database waveforms, output in analogue form, as allowed  under the agreement with Biosigna. However it is still necessary to buy the database from INSERM to access the digital data and reference data.

More information and experience on the CSE database is provided in this 2014 article.

IEC 60601-2-25 Update Guide (2005 to 2011 edition)

For the second edition of this standard, IEC 60601-2-25 and IEC 60601-2-51 were combined and re-published as IEC 60601-2-25:2011 (Edition 2.0).

The standard has of course been updated to fit with IEC 60601-1:2005 (the 3rd edition). Also, similar to IEC 60601-2-27, the opportunity has been taken to correct some of the errors in requirements and test methods for performance tests that existed in the previous edition. However, compared to the update of IEC 60601-2-27, the changes are far more extensive making it difficult to apply the new standard in a gap analysis approach. Experience also indicates that historical data for existing equipment is often of limited quality, so it may anyhow be an excellent opportunity to do a full re-test against the new standard.

Despite the updated tests, it seems that significant errors still persist, which is to be expected given the number of complexity of the tests.

The table below provides an overview of corrections, changes and problems found to date in the new standard. This table was compiled during the real tests against new standard.

One major change worth noting is that requirements for ECG interpretation (the old clause 50.102 in IEC 60601-2-51) have been completely removed from the standard. There is no explanation for this, however the change is of interest for the CB scheme since it is now possible to objectively test compliance with all performance tests.

Table: List of changes, corrections and problems in IEC 60601-2-25:2011
(compared to IEC 60601-2-25:1993/A1:1999 + IEC 60601-2-51:2003)

Clause Subject Type Content
201.1.1 Scope Change

The scope statement has been reworded, so for unusual cases it should be checked carefully.

There has been a common mistake that IEC 60601-2-25/IEC 60601-2-51 should not be applied to patient monitors, and a similar mistake can also expected for this edition. However, the correct interpretation has always been  that if the patient monitor provides an ECG record intended for diagnostic purposes, then diagnostic standard should also be applied.

This would then depend on the intended purpose statement (and contraindications) associated with the patient monitor. However, manufacturers of patient monitors with 12 lead ECG options, with measurements of amplitudes, durations and intervals or automated interpretations might find it difficult to justify a claim of not being for diagnostic purpose.  

201.5.4 Component values Change For test circuits, resistors are now required to be ±1% (previously 2%)
201.6.2 Classification New The ECG applied part must now by Type CF (previously there was no restriction). Detachable lead wires Change Detachable lead wires must be marked at both ends (identifier and/or colour) Instructions for use Change

Requirements for the operation manual have been substantially modified in the new standard (see standard for details).

Note: it seems that HF surgery got mentioned twice in item 6) and 12), possibly as a result of combining two standards (IEC 60601-2-25 and IEC 60601-2-51) Defibrillator proof tests Change

Due to the size of the clause, it is difficult to fully detect all changes. However, at least the following changes have been found:

  • The test combinations (Table 201.103) now includes 12 lead ECGs (i.e. C1 ~ C6 should also be tested)
  • The energy reduction test is now included (previously not required for diagnostic ECGs)
  • The test with ECG electrodes is now removed

The energy reduction test is a major change: many diagnostic ECGs have no series resistors which helps to limit noise, improve CMRR. To pass the energy reduction test, ECG lead wires should have at least 1k resistors and preferably 10k (as discussed in the technical article on Defibrillator Tests). With this series resistance, the impact of the ECG gel electrodes is reduced, and perhaps this is the reason for making the test with electrodes obsolete. The test result anyhow depended on the type of ECG electrodes, which is often outside the control of the manufacturer, making the test somewhat unrepresentative of the real world. Automated interpretation Change Automated interpretation is now removed from the standard. Note that it is still expected to be covered by regulatory requirements, such as Annex X of the MDD. Automated amplitude measurements Correction

The limits stated in the requirements has now been corrected to match the test method (5% or 40uV).

The reference values for CAL and ANE waveforms have now been included in the standard (Annex HH). The previous edition stated that these values were there, but they were missing.


In the Annex HH reference data, the polarity of some S segment values is wrong (CAL 20500, aVL, aVF , and V3 for all of the ANE waveforms). There may be other errors that get revealed with time. Automated interval measurements (CAL/ANE) Problem

The requirement statement refers to global measurements (with 17 waveforms, up to 119 measurements), however the compliance statement refers to measurements from each lead (for a 12 lead ECG, up to 1428 measurements if all durations/intervals are measured). Not all ECGs provide global measurements, so this really should be clarified.

Because of this it is also unclear about the removal of 4 outliers "for each measurement". If global measurements are used, this would imply that 4 out of the 17 measurements can be removed from the statistical analysis (which seems a lot). However, if lead measurements are used, this implies 4 out of 204 measurements, which is more reasonable. 

201.12.4 General test circuit Correction / change

The test circuit is now correctly and harmonized with IEC 60601-2-27:2011, IEC 60601-2-47 and also ANSI/AAMI EC 13. Previously the 300mVdc offset was placed in parallel with the test signal which meant the impedance of the dc supply appeared in parallel with the 100Ω resistor and reduced the test signal. The dc offset is now placed in series where this problem does not occur.

However, it is noted that for one test the 300mV DC offset is still required to be applied "common mode" using the old circuit.

Also, in the old standard the resistance between RL/N to the test circuit was 100Ω, whereas now it is a 51kΩ//47nF. A conservative interpretation is that all tests should be repeated with the new circuit, given the significant change (although experience indicates the results don't change). Indication of inoperable ECG Problem The standard indicates that the test should be performed with 1V steps, up to 5V. However, the point of saturation normally occurs well below 1V (experience indicates this is from 400 - 950mV). This means it is possible to pass the test, without passing the requirement. The standard should instead require the dc voltage to be increased in steps of 5 or 10mV to ensure that the indication of saturation is provided before the signal amplitude starts to reduce. 
Test of network Change The previous test (application of 2mV and 6mV waveforms to various leads) is now replaced with the CAL and ANE waveform, with a limit of 10%

The above change has interesting points. The first is that one might ask why the test is needed, since the CAL and ANE waveforms have already been tested under (automated amplitude measurements). However, Clause can be done by digital analysis, whereas this test is for the full system including the ECG's hardware. Also, not all ECGs measure all amplitudes. 

It therefore requires the ability to generate CAL and ANE test signals by analogue (with laboratory 1% accuracy) which many laboratories may not have.

That said, the test really does not really seem to test the networks correctly. As in the old standard, the networks are best tested by providing a signal to one lead electrode only, whereas the CAL/ANE waveforms provide the same signal to all leads simultaneously, except RA which is grounded. Although some analysis is required it seems clear that at least part of the lead network cannot be tested by the CAL/ANE waveforms.

Finally, one might ask why there is a 10% limit for the test method, while the requirement statement says 5%. The reason could be that the basic measurement function is 5%, while the lead networks add another 5%, thus providing an overall 10% error. This is a clear relaxation on the previous edition, which seems unnecessary given that modern electronics (and software) easily handles both the measurement and network well below the 5% in the old standard. Input Impedance Correction

The previous version of the standard had an allowable limit of 18% (for reduction with 620k in series), but Table 113 incorrectly had an effective 6% limit. The 6% limit could be met at 0.67Hz, but most ECGs failed at 40Hz (the input impedance changes with frequency).

The new standard now corrected this to a limit of 20%, aligned with IEC 60601-2-27.

The requirement to test with a common mode 300mV to RL has been removed. Required GAIN Change / Problem

The previous standard included a test of a 1mV step to verify the sensitivity (mm/mV), with a limit of 5%. This test and the limit are now removed, which means there is no objective measurement to verify that a 1mV step corresponds to 10mm on the ECG record. This may or may not be deliberate: it opens the possibility that manufacturers may use a gain of "10mm/mV" in a nominal sense only, with the actual printed record being scaled to fit the report or screen. The classic 5mm pink background grid also then also scaled to give the appearance of 10mm/mV, even though the true measurement reveals strange values such as 7mm/mV (on a small screen) or 13mm/mV (on a printed record).

Using the definition, "GAIN" is the "ratio of the amplitude of the output signal to the amplitude of the input signal". The text in refers to the amplitude on the ECG record. Putting these together, it seems the literal interpretation is that 1mV input should be 10mm on the ECG record. Also several tests provide the allowable limits in mm (e.g. CMRR test limit is 10mm), if the outputs are scaled this would make little sense.

But in the absence of any criteria (limit), it is all a bit vague. If scaling is allowed, it should be clearly stated, and limited to a suitable range otherwise it can get confusing (e.g. at "10mm/mV", a 1mV indication should be in the range of 8-14mm. The scaling and reference grid should be accurate to within 5%, although in the digital world, this may be not necessary to test beyond a spot check to detect software bugs. Finally all limits in the standard should be converted to mV or uV as appropriate. CMRR Change

The test is the same except that the DC offset is now included in the CMRR test box, and the RL/N lead electrode is no longer required to be switched (test is harmonized with IEC 60601-2-27). Previously, the standard indicated that the dc offset is not required, because it has been tested elsewhere. Filter affecting clinical interpretation New

The standard now requires the ECG report to include an "indication" that the clinical interpretation may be affected by filter settings (if applicable). However, there is no clear statement about what is an acceptable "indication". It could mean text such as "Warning the interpretation: might not be valid due to the use of filters"; on the other hand it could mean just making sure that the filters used are clearly shown on the record, perhaps adjacent to the interpretation (allowing the user to make thier own conclusion).

What makes the issue more confusing is that some manufacturers might apply the filters only to the printed waveforms, while the interpretation is still performed on the unfiltered data (to ensure that the filters don't mess up to the interpretation), or worse, some kind of mixed situation (e.g. only mains hum filter is allowed for interpretation). Notch filter effect (on ANE20000) Change

The allowable limit for ringing in the ST segment has now been increased from 25uV to 50uV.

Test experience indicates that the impact of notch filters for the waveform ANE 20000 on Leads I, II and III, aVR, aVL, aVF is minimal. However, the very large S amplitude on Leads V2 (1.93mV) and V3 (1.2mV) can cause a large amount of ringing in the ST segment, which is probably the reason for the change in limit.

It is possible that previous tests have been limited to inspection of Lead II with the assumption that the ANE20000 waveform is the same for all leads (a mistake which the author has made in the past). In fact, the test should be done with a full 12 lead ECG simulator, with each lead inspected one by one. If the notch filter is applied in hardware (in part or full), the test should be done in analogue form. Baseline, general Change Many of the baseline tests have now been removed, such as temperature drift, stability, writing speed and trace width, presumably because in the modern electronic / digital world these are not worth the effort to test. Most ECGs use a high pass filter and digital sampling, which means there is no real possibility for baseline drift. Channel crosstalk Change

The previous edition mixed up leads and lead electrodes (for example, putting a signal on RA/R results in a signal on Leads I, II, aVx, and Vx) so the criteria never made any sense. In practice the test needed to be adapted.

Fortunately, this test has now been corrected and updated to give clear direction on where lead electrodes should be connected and also which leads to inspect for crosstalk. The test is the same as in IEC 60601-2-27:2011.


In step c) of the compliance test, the standard says to inspect Leads I, II and III, but this appears to be a "cut and paste" typographical mistake. The correct lead is only Lead I (Leads II, III will have a large signal not related to crosstalk). Similarly, in step d) this should be only Lead III. Steps e), f) and g) are all correct. High frequency response Change

For frequency response, previously all tests A to E were applied, in the new standard only tests (A and E) or (A, B, C and D) are required.

Also the limit for test E has been slightly reduced (made stricter) from -12% to -10%. Low frequency response Change

The allowable slope has been changed from 250uV/s to 300uV/s, perhaps in recognition that a single pole 0.05Hz high pass filter (typically used in many ECGs) could not pass the 250uV/s limit. Theoretical simulations showed that 0.05Hz single pole filter produces a slope of 286uV/s.

Problem Minor mistake in the standard: the requirement statement does not include the limit for the slope of 300uV/s. This is however included in the compliance statement. Linearity and dynamic range Change / problem

The previous test method used a 1mVpp signal, but required the minimum gain. For an ECG with typical minimum gain of 2.5mm/mV, this meant that the test signal was only 2.5mm, which then conflicted with the diagram.

The new standard corrected this, but made the slight mistake of saying "10mV" rather than "10mm". But the test only makes sense if 10mm is used. Time and event markers Change / problem

It appears as if the authors of the standard were getting a bit tired by this stage.

Both editions of the standard fail to provide a test method, and it is not really clear what to do. The compliance statement is effectively "X shall be accurate to within 2% of X", which makes no sense.

In the latest edition, things have got worse, with the reference to test conditions referring to a clause that has no test conditions (

In practice one would expect the time markers to be accurate to within 2% compared to either a reference signal (e.g. 1Hz for time makers of 1s), and/or against the printed grid.

Of course, all of this really has not much impact in the digital world with crystal accuracy of 50ppm (software bugs notwithstanding). Pacemaker tests Change

The previous pacemaker tests (51.109.1 and 51.109.2) have been combined and extensively reworked:

  • The requirement statement has been changed to include pacing pulses of 2mV to 250mV and durations 0.5 to 2.0ms
  • The test circuit for pacemaker has been defined
  • the point of measurement of amplitude after the pulse is changed from 50ms to 120ms (3mm)
  • the test with the triangle pulse (or CAL ECGs) is removed
  • the test method now includes a calibration step (item e)) to ensure the 2mV pulse is accurate
  • there is now no requirement to declare the impact of filters
  • (big point) the test is now clearly required for all electrodes, tested one by one as per Table 201.108

Although the requirement statement refers to 0.1ms, there is no test for this. 

Also, the status of filters is not clear. Most ECGs use hardware or software "blanking" when a pacing pulse is detected, leaving a clean ECG signal for further processing including filters. This means that the filter setting has no impact on how the ECG responds. However, some manufacturers don't use this method, allowing the display to be heavily distorted with the pulse, with the distortion varying greatly depending on the filters used. Ideally, the standard should encourage the former approach, but at least if heavy distortion can occur for some filter settings, this should be declared.