Computed tomography “CT” is increasingly becoming a highly reliable imaging modality in medical imaging. Whilst computed tomography techniques had limited use during 70s and 80s, due to the increased scanning speed and improved image quality, CT usage has increased dramatically since then. Today, large numbers of patients are imaged by CT techniques covering almost every part of the body. According to the ICRU, about 250 million CT scans are performed worldwide per year [1]. Even though projection imaging such as chest radiographs remains the dominant radiological imaging modality, CT is recognized as the modality that delivers the highest radiation dose to the public. As an example, according to Kalender et al. [2] and the German Federal Office for Radiation Protection [3] whilst only about 7% of total x-ray imaging in Germany was done by CT, 60% of the total dose to the public from all the x-ray imaging modalities was attributed to the CT scans. It is highly unlikely that Germany is unique in this regard.

Computed tomography is one of the most common imaging modalities for brain and head and neck [4]. For brain imaging CT has many advantages over other imaging modalities including the fact it is fast imaging technique, hence it less affected by patient motion, less expensive, easier to perform, provides detailed images most required in brain imaging, and can be used at no risk to the patient with implantable devices. Though CT acquired images are normally with great contrast, it is required in more than half of the cases a contrast media inclusion in the target prior imaging [4].

Whilst radiation dose delivered during CT examination is an important consideration, it is generally not measured directly but taken into account indirectly through a dose indicator. This is because the distribution of the dose delivered by CT imaging through slices is not uniform and it cannot be measured directly in the patient using known types of dosimeters. Shope et al. [5] introduced the definition of computed tomography dose index/indicator “CTDI” in 1981 as a parameter to indicate the level of doses delivered by CT scanners. Since then many modifications of the CTDI basic form have been introduced to accommodate new conditions and new instrumentations introduced in this field such as the pitch value which is introduced through the spiral CT techniques. The full meaning and description of this parameter “CTDI” is discussed by McCollough et al. [6] and also the references cited in this paper provide a comprehensive review of the changes introduced to it over the years. For example, the dose estimated over volume and the difference between the dose-levels at the centre of the slice as compared to that on the periphery are currently included as corrections. Full details of employing CTDI to estimate dose delivered by a CT scanner with all the corrections can be found in Kalender et al. [2]. The IAEA code of practice for dosimetry in diagnostic radiology “TRS457” recognises the concept of CTDI as defined by the IEC [7], measured using a 100 mm pencil ionization chamber. CTDI is a dose indicator and all of its modified forms such as Dose Length Product “DLP” are also dose indicators. CTDI is defined for sequential CT. For volume scanning based on the helical principle, the pitch is included and the CTDI is then termed volume CTDI, “CTDI_vol”. The DLP is the product of the CTDI_vol and the scan length and it is utilised for multislice slice and multi detector CT scanners. DLP can be converted crudely to measure the effective dose through the use of some conversion factors. Seibert et al. [8] recently introduced the effects of patient size on the CTDI_vol and DLP values. Also recently a comprehensive explanation for the dose delivered by CT has been documented in an AAPM report [9]. Contrast media, such as iodine-based compounds, enhance tissue contrast due to its higher atomic number and physical density compared to the tissues. These characteristics dramatically increase the probability for the photoelectric effects & Compton interactions and this leads to enhancing the CT image contrast [10]. Though CT images show far superior details of the target compared to projection imaging i.e. contrast resolution is higher, in some cases still further details are required clinically; therefore, contrast media of the type traditionally used in projection imaging (iodine compound) is required and used to enhance the image details by increasing the subject contrast. The contrast media is used in CT imaging to improve the visualisation of anatomic structures (usually blood vessels or parenchymal organs) that would be indistinguishable from background tissue, in terms of their x-ray attenuating properties. A phantom based study by Jackson et al. [10] showed the levels of contrast enhancement by CT scan when various types of contrast media are used.

For CT scanners equipped with current modulation, it has been shown that the indirect effect of contrast media (the tube current is increased to compensate for the attenuation of the CM, and the increased current increases the CTDI) is significant enough to change clinical practice [11]. It has also been proven, through radiobiological studies, that inclusion of the contrast media in the tissue of interest (target) directly enhances the dose locally (or radio-sensitises the cells) at the target [12 - 17]. Generation of copious secondary electrons, i.e. free radicals from the interaction of x-ray photons with the higher density and particularly higher atomic number of the iodine atoms, leads to an increase in the photoelectric effect probability of interaction and this is the main reason underlying dose enhancement. However, the direct effect of contrast media on the dose and hence on the CTDI value is currently not factored into the equation for the determination of the CTDI value. The CTDI for newer CT scanners with Tube Current Modulation is affected indirectly by an increase in tube current, and this is an active area of research [11] but currently for all CT scanners the method for calculating the CTDI does not take into account any direct change in dose due to the presence of contrast media. This research is based on the evaluation of the direct effects of the contrast media on the CT delivered doses and particularly its currently unaccounted effects on the CTDI value. The work comprises both theoretical calculations and experimental measurements. Moreover, this contrast factor for dose increase is derived in terms of the average CT number at the target. This means this contribution can easily be estimated from the known CT numbers at the target. The contribution of the contrast media to the CTDI value depends on the concentration of the media in the target and also on the kVp value. This makes it feasible to determine the levels of dose enhancement directly from the CT numbers of the imaged target.

Paul et al. [18] conducted a study on patients CT data to determine the effects of the contrast media on the image noise and the dose delivered. The effects on dose were inferred from the CTDI_vol value changes between conditions of non-contrast with the existence of the contrast media to the same targets. Their results showed an increase in CTDI_vol values between 3 to 13 per cent. A clinical study by Amaton et al. [19] also identified that inclusion of iodine contrast media in CT scans results in direct dose enhancement specific to various organs ranging from 20% to 70%. In addition, a recent Monte Carlo simulation based study conducted by He et al. [20] also demonstrated an increase in dose caused by micro-spheres containing contrast media in a modelled phantom. Whilst these studies indicate (or suggest) that contrast media does have both a direct and indirect impact on CT dose, there is a need to theoretically and experimentally examine and quantify the direct effects of contrast media on CTDI.

In particular, this paper includes theoretical and experimental dose enhancement determinations following the administration of varying concentrations of iodine-based contrast media just prior to scanning, as well as modifications in the CTDI equation. This work also relates dose enhancement of contrast media to the CT number. Hence, to estimate the contrast media effects on the CTDI it is only required to know the CT numbers at the target i.e. slice. The dose enhancement at these levels of low doses are calculated according to the method introduced by Corde et al. [12] for radiotherapy beams which agrees with the dose enhancement values obtained from the CT numbers. Moreover, in this work we also extended the notion of this simplistic method of dose enhancement estimation based on just mono-energetic beams into using spectral energy values of the x-ray photons and under more realistic conditions of low levels of contrast media concentrations (similar to those expected in humans before CT scanning). In this research the dose enhancement and its relationship to the CT numbers is derived from the first principles and related it to the measured values at certain concentrations of the contrast media.

Recently, a review article is documented on the effects of iodinated contrast media on the cells’ DNA in diagnostic x-ray imaging in general [13 , 14]. High dose enhancement levels are reported by these researchers in the intra and extra vascular and in the blood and in the microdosimeteric environment.

Direct dose enhancement dependence on beam energy is an important factor which has attracted the attention of many investigators in the therapy field especially those interested in heavy atom based nanoparticles [22]. The apparent disagreement between the observed therapy experimental data and theoretical predictions needs to be resolved, as all investigations showed that maximum dose enhancements occur at kV ranges of x-ray energies which include the levels of x-ray energies used in CT imaging, yet the direct effect of contrast is ignored in all diagnostic imaging dose calculations..

In this research the dose enhancement is derived from the first principles of energy deposition and then linked to the CT numbers as shown in Appendix 2. Moreover, Equation (5) also relates the dose enhancement to the observed CT number of the target tissue type. This means if the CT numbers for an organ with the contrast media is known it can be used to determine the levels of direct dose enhancement by the contrast agent to the organ. To proceed from the DEF to an organ to the DEF for the patient CTDI, one either has to know the weights for each organ and relative dose to each organ, or more practically in the absence of such data, to assume uniform dose to each organ.

The dose enhancement is attributed to the higher absorption rate of x-rays by iodine atoms in the contrast media leading to an increase in the generation of secondary electrons, i.e. free radicals including Auger electrons. More details on the principles of dose enhancement caused by iodine based contrast agents and metallic nanoparticles and gold in particular can be found in the literature [10 , 22].

The direct dose enhancements caused by inclusion of set concentrations of contrast media in phantom model CT targets are displayed in Figure 1 & Figure 2. For instance the dose enhancement as determined by the CT number changes due to the inclusion of the contrast media from Table 2 at 20% concentration is about 16% as obtained using Equation (5) which relates the levels of dose enhancement to the CT number.

For standard CT clinical protocols, however, it is possible to show that the DEF for uniformly distributed contrast within the scanned volume is not more than 0.6%. The assumptions in making this estimate are crude: that the patient is a uniform cylinder, that the contrast is uniformly distributed within the irradiated volume only, that all of it is present at the time of the CT scan, but the model has the virtue of using only information generally retrospectively available for review, and the relaxation of any of these assumptions would reduce the DEF. If some of the contrast is present outside the irradiated volume, for example, the DEF will reduce. The model is so crude it should not be used to compare clinical protocols, as the model implies that reducing the scanned length increases the DEF. The only exception to this is the postulate of uniform contrast distribution. In clinical cases, contrast is not uniformly distributed; it is unevenly distributed in a way that enhances the clinical usefulness of the image. It may therefore be concentrated mainly in a relatively radiobiologically sensitive organ, and in this circumstance, it is conceivable that the presence of contrast may produce a significant DEF, as has been noted in the literature. For the case of studies with a high concentration of contrast in a radiosensitive organ, the DEF in the organ can be estimated from the CT number change due to the contrast (using the values in this work as an approximation) and the radiobiological weight factor of the organ.

In contrast, the possibility of a rescan due to a failure to inject contrast is potentially a much larger factor for individuals, but clinical cases of this are handled by legislated error reporting systems, and in studies the ethics approval often requires that studies with missed contrast are not repeated and the patient is excluded from the study.

The indirect effects of contrast for CT scanners with Tube current modulation are potentially much higher, but they are adequately taken into account by the standard CTDI equation, and so they are not a major focus in this work.

Administering contrast media to the CT patients is mainly to enhance the image contrast. However, it has been long proven experimentally that it also enhances radiation dose at the target depending on its concentration at the target. Infliction of radiation dose to patients individually and to the public in general has also been the focus of large number of investigations as documented in the literature. This research quantified this dose enhancement and introduced it into the CTDI indicator. Moreover, this work also shows a simple and direct method of estimating the dose enhancement caused by contrast media based on the CT number change due to its introduction.

The dose enhancement caused by the introduction of the contrast media into the radiation target is concentration at the target dependent. At clinically administered concentrations of contrast media, a measurable dose increase is obtained justifying its inclusion into the CTDI equation. In so doing, the level of dose enhancement inflicted by the contrast media can be estimated from the CT number change at the target in the CT slice.

However, it should be noted that inclusion of contrast in a CT scan dose equation i.e. the CTDI does not quantifiably increase the radiation dose burden to the patient unless there is a high concentration of the contrast aggregated in a radiosensitive organ such as breast or brain.

We thank the support of William Buckland Radiotherapy centre at Alfred Hospital in Victoria Australia for making available the radiation facilities for this work.

The authors declare no conflicts of interest regarding the publication of this paper.

The basic definition of the dose enhancement factor “DEF” as a function of concentration “c” and energy “E” is as follows

DEF ( c , E ) = c I [ μ e n ρ ] E I + ( 1 − c I ) [ μ e n ρ ] E H 2 O [ μ e n ρ ] E H 2 O

Here ω represents the concentration of the solvent [in this case Iodine CM] and

μ e n / ρ is the mass-energy absorption coefficient

However, [ μ e n / ρ ] is mass-energy absorption coefficient which is related to the linear attenuation coefficient in the following way {The Physics of Radiation Therapy Chapter 3 section 5-4B by Faiz M. Khan 5^th edition 2015}

μ e n = μ t r ( 1 − g )

where µ_tr is the transfer energy factor and it is directly related to the linear attenuation coefficient µ as follows;

μ t r = ( E ¯ t r / h ν ) ⋅ μ

where E_tr represents average energy transfer

Therefore

μ e n = ( E ¯ t r / h ν ) ⋅ μ ( 1 − g )

Assumptions:

· Since transfer energy “E_tr” will be almost the same for both cases of tissue equivalent material with and without the iodine contrast agent.

· The “g” factor representing the energy generated inside the target and deposited outside which is not balanced by the electron equilibrium and assuming that the g factor is also the same in both cases. Moreover, radiative loss factor is negligible in case of the beam qualities typically used in CT.

· Also, the density “ρ” is considered to be un-effected by the inclusion of small amount of iodinated compound. Otherwise the final relationship will still be the same however with another factor related to this difference. The CTDI factor is a dose indicator based on many assumptions and approximations which justifies neglecting this difference. Therefore

μ e n = Y ⋅ μ

where Y = ( E ¯ t r / h ν ) ⋅ ( 1 − g )

Inserting the above relationship in the DEF equation will be;

DEF ( c , E ) = ω I ⋅ Y ⋅ μ I + ( 1 − ω I ) Y ⋅ μ ω Y ⋅ μ ω

This can be re-written as;

DEF ( c , E ) = ω I ⋅ μ I + ( 1 − ω I ) ⋅ μ ω μ ω

and can further re-written as;

DEF ( c , E ) = ω ( μ I − μ w ) μ w + 1

But the fraction in the above equation is nothing but CT#/1000

Call CT# = N

Hence

DEF ( c , E ) ≈ ω I ⋅ N / 1000 + 1

In case of no added contrast media the CT number should be that of the target. For instance if we deal with water then CT # i.e. N= 0 and there will be no dose enhancements and its value will become one.