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CT Parameters: How Each Setting Changes Image and Dose

kVp, mAs, pitch, collimation, kernel… every parameter on the CT console shifts the balance between image quality and dose. We cover them all, grounded page-by-page in our reference text, Bushberg.

Intermediate· ~16 min read· Sourced
Scope of this article
All technical content is grounded in our reference text, Bushberg, The Essential Physics of Medical Imaging (3rd ed., 2011), with citations to the relevant pages.1

How the image forms

In CT, the X-ray tube and the opposing detector array rotate around the patient; the attenuation profiles (projections) acquired from each angle are reconstructed by computer into a slice image. In multidetector (MDCT) systems, several slices are acquired per rotation, and in helical (spiral) mode the table advances through the gantry while data are acquired continuously.1 The real skill is knowing how each parameter shifts the balance between image quality and dose.

kVp — tube voltage

kVp sets the energy (penetration) of the beam. Higher kVp gives a more penetrating, "harder" beam; it lowers noise and ensures sufficient photon transmission in large patients, but reduces tissue contrast (especially iodine contrast). Lower kVp raises the photoelectric effect of iodine and thus increases contrast — which is why low kVp (e.g. 80–100 kVp) in slim patients both boosts contrast and lowers dose in angiographic/contrast studies. Modern scanners include automatic kV selection that suggests the optimal voltage per exam.1

mAs and tube current modulation

mAs = tube current (mA) × rotation time. It directly sets the number of photons, hence noise and dose. Doubling mAs roughly doubles dose. Image noise, however, is inversely proportional not to dose but to the square root of dose: halving noise requires roughly four times the dose (mAs).1

Automatic tube current modulation (ATCM)
Modern scanners vary mA in real time according to the patient's anatomy, both along the z-axis and by gantry angle (in an approximately sinusoidal pattern, vendor-dependent), giving more uniform noise across the image and avoiding unnecessary dose.1

Pitch

In helical scanning, pitch is the table advancement per full gantry rotation divided by the total beam width (nT):

pitch = table feed (per rotation) ÷ total beam width (nT)

Per Bushberg, for most CT scanning pitch ranges between 0.75 and 1.5; pitch = 1.0 corresponds in principle to contiguous axial scanning. Below 1.0 (overlapping) patient dose rises; above 1.0 it falls. With all else equal, the relation between relative dose and pitch is:1

relative dose ∝ 1 / pitch

In practice: high pitch (≈1.5) speeds up thoracic and pediatric scans and lowers dose; low pitch is used where high temporal/spatial consistency is needed, such as cardiac.1

Collimation, slice thickness and detector configuration

In MDCT, the total beam width (nT) is set by the collimator and the slice thickness by the width of the detector arrays (T). For example, a 64 × 0.5 mm system can re-reconstruct different slices from the raw data — 64 × 0.5 mm, 32 × 1.0 mm, 16 × 2.0 mm, etc.1 Thinner slices improve spatial resolution but raise noise, since fewer photons contribute per slice.

Overbeaming and geometric efficiency
In helical MDCT the beam penumbra is placed outside the active detector array (overbeaming), creating a dose inefficiency. Early systems with few arrays had low geometric efficiency (≈70%, ~30% dose penalty); modern 64-channel systems exceed **95%**.1

Reconstruction kernel (filter)

Different kernels produce different images from the same raw data. A sharp kernel gives high spatial resolution (for bone, lung) but more noise; a smooth kernel lowers noise (for soft tissue, brain) but reduces sharpness. Bushberg emphasizes the kernel's decisive effect on the modulation transfer function (MTF) — hence on resolution.1 Field of view (FOV) and matrix size also set pixel size and thus resolution.

Noise, contrast and resolution

Image quality has three axes, and parameters pull them in different directions:

The essence of optimization is to reach the quality needed for diagnosis at the lowest dose; "a better image" does not always mean more dose.

Summary: which parameter changes what?

Parameter Increasing it Image effect Dose effect
kVp Beam hardens Noise ↓, contrast ↓ Dose ↑ (at constant mAs)*
mAs More photons Noise ↓ Dose ↑ (linear)
Pitch Faster table Possible z-axis resolution ↓ Dose ↓ (∝ 1/pitch)*
Thin slice Smaller voxel Spatial resolution ↑ Noise ↑
Sharp kernel MTF ↑ Resolution ↑, noise ↑ Dose unchanged
Small FOV Smaller pixel Resolution ↑ Dose unchanged

*These dose arrows hold with all other parameters constant. At constant mAs, raising kVp increases dose; but with automatic kV selection and tube-current modulation (ATCM), the real dose and image-quality effect varies with protocol, patient size and vendor.

International DRL examples

Diagnostic reference levels (DRLs) are an optimization tool, not a limit; they are not used to judge a single patient's dose, but show the typical dose level of a typical patient group (usually the 75th percentile). In CT the DRL quantities are CTDIvol and DLP, using a 16 cm phantom for head and a 32 cm phantom for body.2

Below are examples of adult CT DRLs from two countries (CTDIvol: mGy; DLP: mGy·cm). UK values are published by clinical indication, whereas the Australian values are anatomical exam categories, so the table is not a strict like-for-like clinical comparison:

Examination UK CTDIvol UK DLP Australia CTDIvol Australia DLP
Head 47 790 45 820
Chest 8.5 290 8 310
Abdomen–pelvis 10 530 10 480

UK values are taken from the UKHSA adult CT NDRL table: acute stroke for head, lung cancer for chest, and abscess for abdomen–pelvis.3 Australian values are from ARPANSA's updated adult-CT DRLs.4

Differences between countries can be large
The DRL for an apparently similar exam can vary by country, clinical indication and definition method, because each national DRL is based on its own patient population, scanner fleet and data-collection approach. So a facility's typical (median) dose should be compared against its own national DRL in the same clinical context; if it approaches or exceeds it, the protocol is reviewed.2

The current trend is toward clinical-indication–based DRLs rather than anatomical regions: Europe's EUCLID project (2021) defines DRLs by indication such as stroke, pulmonary embolism and appendicitis;5 and Japan's national DRL set published by J-RIME in 2025 is one recent national example, showing a downward trend in CTDIvol/DLP versus earlier surveys.6

Related articles
For the definitions of the dose metrics: Understanding CTDIvol, DLP & SSDE. For reconstruction generations: FBP, Iterative & Deep Learning. Key terms: CTDI · DLP · Pitch · Kernel.

References

  1. Bushberg JT, Seibert JA, Leidholdt EM, Boone JM. The Essential Physics of Medical Imaging, 3rd ed. Lippincott Williams & Wilkins, 2011. Bölüm 10 (Computed Tomography) ve Bölüm 11 (X-ray Dosimetry). Atıflarda belirtilen sayfa numaraları bu baskıya aittir.
  2. ICRP Publication 135. Diagnostic Reference Levels in Medical Imaging. Ann. ICRP 46(1), 2017. icrp.org
  3. Public Health England (UKHSA). National Diagnostic Reference Levels (NDRLs) — bilgisayarlı tomografi (güncel ulusal değerler). gov.uk
  4. Lee C, Goergen S, et al. Updated Australian diagnostic reference levels for adult CT. J Med Radiat Sci 2020 (ARPANSA). PMC7063242
  5. Damilakis J, et al. European diagnostic reference levels and typical doses for CT based on clinical indications (EUCLID). Eur Radiol 2021. Eur Radiol
  6. J-RIME. National Diagnostic Reference Levels in Japan (2025). j-rime.qst.go.jp
Note: This content is for education; for clinical decisions or regulatory compliance, consult a qualified medical physicist and current regulations.

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