Wearables in Dermatology: Quiz on Next-Generation Technical Advances

ANSWERS AND EXPLANATIONS

1. c

Explanation: The epidermis, particularly the stratum corneum, has very high electrical resistance due to its keratinised, low-water-content cells. This barrier property limits conductivity and complicates the acquisition of biopotentials [e.g., Electrocardiogram (ECG) and Electromyogram (EMG)] through wearable devices.^[1]

2. b

Explanation: TENGs operate based on the combined effects of triboelectrification and electrostatic induction. Triboelectrification occurs when two dissimilar materials come into frictional contact, such as through skin movement; one material donates electrons while the other gains them, generating opposite surface charges. Electrostatic induction arises when these charged materials are separated, creating an electric potential difference that drives electrons through an external circuit to balance the charge distribution.^[2,3] The self-powered nature of TENGs, along with their high sensitivity, durability, flexibility and comfort, makes them highly suitable for wearable skin devices.

3. b

Explanation: Successful attachment of sensors to the skin requires selecting a material with mechanical properties that closely match those of human skin. Silicone-based elastomers, such as PDMS, are commonly used as sensor substrates; however, with a Young’s modulus of approximately 3 MPa, PDMS is considerably stiffer than human skin, which can result in delamination.^[4] In contrast, softer elastomers like Ecoflex (Smooth-On), with a Young’s modulus of around 125 kPa, more closely approximate the compliance of skin, allowing for better conformal contact and stable sensor adhesion.^[5]

4. b

Explanation: The major hurdle for long-term wearable sensor adhesion to the skin is the natural regeneration cycle of the stratum corneum, the outermost layer of the epidermis. The keratinocyte cycle (which includes the renewal of the stratum corneum) lasts between 28 and 50 days, depending on age. This continuous shedding and renewal process causes the outermost layer of the skin to slough off, leading to the detachment of sensors adhered to the skin’s surface.^[6]

5. d

Explanation: The dermis contains dense collagen bundles and fibrils, which cause Mie and Rayleigh scattering of light, distorting optical signals.^[7] Scattering reduces the penetration depth and signal-to-noise ratio of reflected/absorbed light, complicating measurements like blood oxygenation or glucose levels.

6. b

Explanation: Microneedles produce almost no pain or very little pain because they penetrate up to where no large ending nerves are present and are minimally invasive.^[8]

7. b

Explanation: Researchers at the Department of Dermatology and Allergy, Technical University of Munich, are developing a miniaturised wristband sensor designed to provide personalised predictions for atopic eczema. This device measures EDA, which reflects the skin’s electrical resistance and is influenced by the sympathetic nervous system and sweat gland activity, thereby correlating with stress levels. Previous studies have explored the utility of EDA as a diagnostic marker in other medical fields, including perioperative stress monitoring in surgical specialties and the assessment of severe depression in psychiatry.^[9,10]

8. b

Explanation: Wearables provide longitudinal at-home data, supplementing clinic-based evaluations without replacing physicians. It provides continuous data for a longer period, because of which we can objectively assess the pattern.^[11]

9. b

Explanation: Pruritus, or itch, is the primary symptom of AD and the main contributor to the reduced quality of life experienced by patients. Since the early 2000s, researchers have sought objective methods to measure scratching, employing technologies such as accelerometer-equipped watches, forearm electromyography, piezoelectric sensors on fingernails, sound recording devices and electromagnetic motion detection.^[12] Among these, actigraphy has emerged as one of the most successful approaches, with recent technological advances making it precise, accurate and reliable for clinical use. A wrist actigraph, a portable device incorporating an accelerometer, measures wrist movements as a proxy for scratching.^[13] Studies have shown that wrist activity between 1 and 3 Hz during the first three hours of sleep serves as a meaningful indicator of AD severity in children.^[14] In addition to quantifying scratching, actigraphy provides valuable data on sleep disturbances in patients with AD.^[15]

10. a

Explanation: Smartwatch-based designs for measuring itch have certain limitations. They are primarily useful for monitoring nocturnal symptoms, as daytime movements can confound the measurements, and they cannot detect scratching performed with the arm not wearing the device – a limitation that is particularly relevant when eczema affects the wrist or arm wearing the watch.^[15] Despite these constraints, smartwatches have demonstrated the ability to identify nighttime scratching behaviour with over 90% accuracy.^[16]

11. c

Explanation: While wearable devices can measure physical parameters such as TEWL and skin hydration, they are unable to assess systemic biomarkers like IgE.^[17] To provide an objective evaluation of AD severity, epidermal probes – such as the Tewameter TM210 and Corneometer CM825 – have been employed. These devices measure TEWL to assess skin barrier function and the capacitance of the skin surface to evaluate stratum corneum hydration. Utilising such noninvasive measurements, the Objective Severity Assessment of AD score has been developed and demonstrated a strong correlation with the SCORAD index commonly used in clinical practice.^[18]

12. a

Explanation: Before wearable sensors can be widely adopted in clinical practice, it is essential to validate their readouts against established measures of disease severity. In the context of AD, this initial validation will likely involve cross-referencing sensor data with current clinical ‘gold standards’, such as the EASI or regional EASI scores.^[17]

13. b

Explanation: Reported challenges with subcutaneous sensors include their relatively short functional lifetime and difficulties with calibration. Another significant limitation is non-specific protein adsorption at the sensor–tissue interface when exposed to complex biological media, a process known as biofouling, which can reduce sensor performance. Currently, subcutaneously implanted sensors are primarily used for glucose monitoring, which may hold particular relevance for dermatological conditions associated with impaired glucose metabolism, such as psoriasis, acanthosis nigricans and hormonal acne. However, data on the use of subcutaneous sensors in skin diseases remain limited.^[19,20]

14. b

Explanation: A novel approach to measuring stress in the skin has been developed through a collaboration between research groups at Stanford University and the Karolinska Institute. They designed a sensor incorporating a specialised membrane capable of non-invasive detection of cortisol in sweat. In the absence of cortisol, the membrane remains permeable to ions; however, when cortisol is present, it blocks ion passage, which is subsequently detected by the sensor. The device demonstrated high reliability, with measurements unaffected by body temperature fluctuations, and showed specificity by not responding to structural analogues such as cortisone. Validation studies using both artificial and human sweat further confirmed its accuracy. Objective measurement of cortisol levels through such sensors offers valuable insight into patient stress levels.^[21]

15. c

Explanation: N-doped carbon dots function as photo-activated nanozymes, facilitating the catalytic oxidation of TMB across all three UV ranges (UVA, UVB and UVC), thereby overcoming the limitation of conventional systems that respond only to UVB and UVC. Importantly, they retain catalytic activity at near-neutral pH (approximately 6), ensuring safety and biocompatibility for wearable dermatological applications. This broad-spectrum responsiveness enables more comprehensive monitoring of UV exposure, which is particularly relevant as UVA constitutes the majority of solar UV radiation reaching the Earth’s surface.^[22]

16. c

Explanation: Graphene-based sensors represent a state-ofthe-art innovation for wearable skin hydration monitoring, owing to their lightweight structure, high flexibility and exceptional sensitivity to subtle physiological changes. Their ability to conform seamlessly to the skin’s surface without causing discomfort allows continuous, real-time and noninvasive detection of minute variations in hydration and elasticity, making them particularly valuable for long-term dermatological monitoring.^[23]

17. b

Explanation: The Corneometer (CM 825, Courage + Khazaka) is widely used to assess stratum corneum hydration by measuring skin capacitance. Results are expressed in arbitrary units, with the final value typically obtained by averaging three consecutive measurements at the same site. This device provides an objective, non-invasive and rapid assessment of skin hydration levels. However, as it is not a wearable technology, its utility is limited to point-of-care or research settings rather than continuous monitoring.^[24]

18. a

Explanation: AD is characterised by immune-mediated inflammation and epidermal barrier dysfunction. Recent advances have demonstrated that key cytokines such as IL-4 and IL-13 can be measured directly in lesional skin using biodegradable hyaluronic acid–loaded microneedle patches, with changes in cytokine levels broadly correlating with clinical improvement. This approach highlights the potential of microneedle-based ISF sampling for detecting inflammatory cytokines, offering a promising avenue for objective monitoring of AD.^[23,24]

19. c

Explanation: Age is often cited as a barrier to the use of wearable devices, as elderly patients may face challenges in adopting and consistently using new technologies.^[25,26]

20. c

Explanation: The image depicts a flexible piezoelectric sensor that generates an electrical signal in response to mechanical deformation such as bending, stretching or applied pressure. Utilising the piezoelectric effect, these sensors convert mechanical strain into electrical voltage and are commonly employed for monitoring arterial pulses, finger movements and respiration patterns. Unlike triboelectric sensors, which depend on friction between dissimilar materials and produce voltage through contact–separation cycles, piezoelectric sensors generate continuous signals under sustained pressure. In contrast, microneedle sensors penetrate the skin to access ISF, while optical sensors operate through light absorption, transmission or reflection, making them non–deformation-based modalities.^[6]

21. b

Explanation: A prospective randomised trial in elderly individuals with actinic keratoses demonstrated a 95% reduction in the incidence of non-melanoma skin cancers among those using UV dosimeters. These devices provided real-time feedback on UV exposure and delivered behavioural cues that encouraged protective measures such as timely sunscreen application.^[27,28]

Close

Figure 5:

Ultraviolet wearables with smartphone apps that provide real-time exposure alerts.

Export to PPT

22. b

Explanation: Many modern UV wearables are integrated with smartphone applications that deliver real-time exposure alerts [Figure 5] and personalised recommendations, such as reminders to reapply sunscreen, hydrate or seek shade. By providing tailored feedback, these devices enhance user engagement and improve adherence to sun safety behaviours.^[23,.29]

23. a

Explanation: Smart bandages incorporating continuous temperature and pH sensors have been tested for monitoring the healing of chronic wounds. In one design, the sensor was integrated into a closed-loop system that released drugs through a thermally responsive carrier embedded within a hydrogel patch whenever real-time measurements indicated inflammation or infection.^[17,27]

24. b

Explanation: Actigraphy captures accelerometer-based movements. Peaks in the 1–3 Hz range correspond to the rhythmic frequency of scratching during sleep in AD patients. Unlike random limb movements, these repetitive bouts correlate with disease severity and sleep disturbance, and they have been validated against clinical scores (EASI, SCORAD).^[16]

25. b

Explanation: TENGs operate through repeated contact and separation of two dissimilar materials, generating surface charges through triboelectrification and producing current through electrostatic induction, whereas piezoelectric devices rely on a single material that generates charge directly when mechanically stressed, producing voltage. Both mechanisms are exploited in dermatology wearables for self-powered strain sensing, motion detection and skin-mounted electronics. Clinically, TENGs are particularly attractive for ultra-light, flexible patches, while piezoelectric sensors are widely used for monitoring physiological signals such as pulse and respiration.^[2,3]

26. b

Explanation: The image depicts a smart, flexible wound dressing designed for real-time monitoring of the wound environment and on-demand antibiotic delivery. This bandage integrates pH and temperature sensors to detect early signs of infection, typically indicated by elevated pH and temperature levels. Upon detection of abnormal values, an embedded microheater activates thermo-responsive drug carriers within a hydrogel layer, releasing antibiotics – such as cefazolin – directly into the wound. These carriers, composed of PNIPAM particles, undergo a physical transformation in response to heat, enabling precise drug delivery. The system operates in a closed-loop manner and supports wireless communication, facilitating remote monitoring and individualised treatment protocols. Such intelligent design makes the bandage particularly suitable for managing chronic wounds, where traditional dressings cannot provide dynamic feedback or controlled drug release.^[27]

27. c

Explanation: The advanced electronic skin developed for stress monitoring simultaneously measures pulse waveform, galvanic skin response and temperature, along with six key sweat biomarkers: Glucose, lactate, uric acid, sodium, potassium and ammonium. Notably, cortisol – a widely recognised stress hormone – is not included among the monitored markers in this system.^[30]

28. c

Explanation: The stress-monitoring electronic skin employs machine learning to analyse multimodal data from physiological signals and sweat biomarkers. This enables the device to distinguish between different types of stress with high accuracy (98%) and predict psychological stress with 98.7% confidence, offering the potential for personalised feedback and interventions.^[31,32]

29. c

Explanation: Waterproof epidermal microfluidic devices are engineered to operate effectively in aquatic environments, such as swimming, or during high-sweat activities. They utilise soft elastomeric materials, including styrenic block copolymers, which prevent contamination from external water while enabling accurate sweat sampling. Embedded colorimetric assays allow real-time quantification of sweat biomarkers, such as chloride concentration, facilitating precise hydration assessment even under challenging conditions.^[33]

30. c

Explanation: A significant barrier to clinical adoption of wearable skin health technologies is the absence of standardised frameworks for interpreting continuous sensor data. Variations in device hardware, software and output hinder uniform clinical application, limiting diagnostic consistency and integration into routine dermatological practice.^[23]