## 1 Introduction  
Thermography is a well-established technology with multiple applications ranging from astronomy to military to construction and building maintenance. It gradually penetrates the healthcare space, including public health and clinical applications. Temperature screening for fever is by far the most well-known healthcare-related application of thermography. However, the utility of medical thermography goes far beyond this application. Core body temperature and skin temperature are two important concepts in medical thermography. Core body temperature refers to the temperature of the body's internal organs, such as the heart, liver, brain, and blood. The average normal core body temperature is generally accepted as being 37 \({}^{\circ}\)C. The core body temperature is tightly regulated by peripheral and central receptors integrated within the hypothalamus [1]. In contrast to the core temperature, the skin temperature changes with the temperature of the surroundings. The skin is the body's largest organ, covering approximately 2 sq\(\cdot\)m, which plays a vital role in thermoregulation. Together with adipose tissues, it provides thermal insulation to the body. In addition, it functions as a \"heat radiator\" system, which is responsible for approximately 90% of body heat loss. Even though public health thermographic screening is based on measuring the core body temperature (or its proxy), most clinical applications of thermography exploit the skin temperature to make conclusions about local tissue physiology or pathology. The clinical applications of thermography in surgery, wound care, and sport medicine were recently reviewed in [2]. The literature review identified two primary use scenarios for thermographic imaging: inflammation-based and perfusion-based. The blood flow to the skin is the most effective mechanism for heat transfer from the body core to the environment. Therefore, the skin temperature is subject to changes and is usually lower than the core [3]. The balance of the thermal exchange between internal and external environments is sometimes described by a so-called _Delta_, which is the difference between the core temperature and skin temperature. In normal conditions, the _Delta_ is around 3.6 \({}^{\circ}\)C (normal core temp is 36.6 \({}^{\circ}\)C, normal skin temperature is 33 \({}^{\circ}\)C). The deviations from this norm can be indicative of inflammation (elevated local temperature) or reduced perfusion (decreased local temperature). While the medical applications of thermography have been known at least since the 1930s [4], its widespread adoption in routine medical practice beyond clinical research and thermographic screening for fever has not started yet. Two primary factors explain it: (a) the technology is expensive for routine clinical use, and (b) there are no simple clinical algorithms. The cost barrier is about to change with an influx of inexpensive uncooled long-wave infrared (LWIR) bolometers. While multiple evidence-based criteria have been proposed on the clinical algorithm side, there are no analytical explanations for these empirical rules of thumb [5]. Thus, building thermophysical models is important in driving clinical technology adoption. This article aimed to elucidate and quantify the primary factors which impact skin temperature. Some of them (e.g., blood perfusion and ambient temperature) are well known [6]. Thus, the primary aim of this article was to identify and quantify other relevant factors. While multiple skin heat transfer models have been proposed, they mostly use either numerical methods [7; 8; 9; 10] or Monte Carlo [11] simulations. Even though these approaches are the only way to solve the transient bioheat equations for complex geometries, some situations (namely thermal acclimation) can be approximated well be stationary solutions. Fortunately, thermal acclimation is a standard for passive thermographic measurements at healthcare facilities.