Bone is an important tissue in the human body as it allows vital locomotive functions of humans to take place, such as providing stiffness, walking or jumping [1]. However, bones are also load-sensitive and semi-brittle [2, 3] and, therefore, prone to injuries. As a result, bone surgical interventions are a daily task in orthopaedics, neurosurgery, and dentistry. However, many of these involve the mechanical material removal by principally cutting (i.e., milling, sawing, drilling) [4, 5] or abrading the tissue to perform the surgery (e.g., total knee arthroplasty, hole drilling, implant or screw insertion). Contrary to soft tissues which are largely elastic [5], bone is semi-brittle, implying that it can easily get damage induced to its subsurface [6] due to mechanical machining [7]; but bone also possess a unique self-healing capacity [8]. These two characteristics make bone an interesting and complex material from both an engineering and medical perspective, which makes tool design a difficult task [9, 10], but it can be better analysed by understanding its intricate hierarchical microstructure [11].

Bone can be broadly divided in two main types, which are the cortical tissue and the trabecular tissue, both of which are shown as a portion of a long bone structure shown in Figure 1. Out of these two, cortical bone is of vast interest since it possesses a superior mechanical strength and density over trabecular bone, thus it is of major relevance when it comes to machining. Cortical bone can be regarded as a natural biological composite material comprised of osteons (i.e., the equivalent to unidirectional fibres in conventional composites) immersed in the interstitial lamellae (i.e., the matrix in conventional composites), as shown in Figure 1. Additionally, the Volkmann’s and Haversian canals, which carry the necessary nutrients for a proper bone function [12], provide a unique arrangement of a porosity system within the tissue. This is one of the reasons why scaffold and implant design is tailored or customised with a controlled porosity that resembles that of natural cortical bone [11], especially if employing additive manufacturing techniques for fabrication [13, 14].

Figure 1
figure 1

Schematic overview of the bone structure: (a) a long bone possesses two main tissues, i.e., cortical bone in the outermost region and trabecular bone in the innermost region, which encloses the yellow bone marrow, (b) the cortical bone microstructure consists of osteons that run along the principal loading direction of the bone and are immersed in the interstitial lamellae. Inside the cortical tissue two main porosity systems (Haversian and Volkmann’s) carry the necessary blood vessels and nerves for maintaining a healthy tissue. The osteocytes (shown as purple dots) are the cells that monitor the metabolism and remodelling process of the tissue. Schematic adapted from Ref. [11]

As a biological tissue, bone possesses a cellular structure that maintains it in its living state with appropriate metabolism and remodelling (i.e., capacity to heal itself over time) functions. One of the cells involved in this process are osteocytes, which are trapped inside the bone in pits called lacunae (see Figure 1(b)) [15]. These cells, among others, keep the bone in a healthy state and aid in the healing process after the tissue is exposed to damage (e.g., by machining-induced mechanical and/or thermal loads [7]).

The intricate microstructure in cortical bone results in a complex material from a machining perspective that makes bone a difficult material to cut, as it can behave both ductile or brittle, depending on the cutting parameters [2]. Thus, during conventional machining, the mechanical effect may induce sufficient strain that facilitates crack propagation from the cutting zone and into the material [16], or induce a brittle and weak layer in the machined subsurface at the micro-level [7]. Additionally, the cutting energy will also dissipate in the form of heat, and as a biological material, the cellular integrity of the tissue may be compromised if the temperature is high enough [17]. In general, if the tissue is exposed to a temperature of 47 or 50 °C during 60 or 30 seconds, respectively, necrosis (i.e., death of osteocytes) will occur, which implies that the healing, remodelling and metabolism functions will be hindered either temporarily or permanently [18]. Additionally, much less time is required for necrosis to occur if the temperatures are larger.

Since mechanical machining (e.g., milling, grinding, sawing) poses a great challenge in terms of machining-induced damage avoidance, non-conventional machining techniques could represent a promising approach towards improving the material response during machining. Laser (e.g., pulsed or continuous lasers) machining can remove material through a range of mechanisms. Pulsed laser ablation is a technique that removes material through the use of a high intensity, focused laser beam, with a short pulse width [19] which leads to energy absorption by the material, subsequently vaporising/ionising it. On the other hand, lasers with longer or continuous pulses may have fluencies below the ablation threshold—leading to a more thermally based material removal mechanism such as melting [20].

Additionally, laser machining provides an improved cutting geometry and precision over conventional tools due to reduced/no mechanical loads [21]. Therefore, laser ablation represents a non-conventional machining approach that could improve the bone cutting process by eliminating the mechanical effect. However, the thermal and shock waves from the laser beam can propagate deep into the bone, thus resulting in possible damage even in regions far from the machined or ablated zone [22]. Due to the fully thermal phenomenon and considering the low thermal conductivity of bone [23], necrosis becomes more relevant.

The three most common lasers used in hard tissues, like cortical bone, are CO2, Er:YAG and Nd:YAG types. The reason for these relies on the optical absorption properties of the tissue, and more specifically, on the ones from its individual constituents (e.g., water, collagen) [24]. For instance, the CO2 laser wavelength (i.e., 10.6 µm) is within the optical absorption spectrum of hydroxyapatite (i.e., 9–11 µm) [25], one of the most important constituents from the bone’s mineral, being the mineral phase the main constituent of bone [26]. The wavelength of Er:YAG lasers (i.e., 2.94 µm) is similar to one of the largest absorption peaks of water (ca. 3 µm), the third main constituent of bone [26]. The wavelength from Nd:YAG lasers is 1.094 µm, but this value can be shortened by frequency doubling, meaning that it can be tailored to minimise water absorption [27], thus maximising the laser penetration depth into the tissue in water-rich (e.g., with coolant) environments.

Rayan et al. [28] did in vivo CO2 laser cutting without coolants in cortical bone and reported that tissue carbonisation is easily induced along with necrosis; however, the laser pulse width was kept constant at 0.1 ms along with a frequency of 2 kHz. Krause et al. [25], also using a CO2 laser without coolants, reported a necrotic depth ranging from 30 to 200 µm when the laser energy density is within 160–2062 J/cm2; however, a comprehensive assessment is hindered since the pulse width (5 ms, 10 ms and continuous) and frequency (10 Hz, 20 Hz, continuous) varied for each energy density condition. Frentzen at al. [29] also employed a CO2 laser (80 µs pulse width) in ex-vivo bone cutting, but due to the usage of external coolant (fine air-water spray), both necrosis and tissue carbonisation were totally avoided.

Since laser machining implies a lower material removal rate as opposed to conventional machining, studies on improving the cutting efficiency in cortical bone have also been explored. For instance, a research using an Er:YAG laser (300 µs pulse width and 20 Hz frequency) [30] proposed a set of laser processing parameters that could result in lower cutting time, when compared to conventional drilling, even producing less heat. However, the laser was coupled with a water spray cooling system and the conventional drilling parameters were intentionally produced with a low speed, thereby hindering the understanding of the real material response both in terms of necrosis and cutting efficiency. Baek et al. [31] showed that mechanical machining, often resulting in a smear layer in the cut surface, blocks the bleeding of the bone, whereas laser machining eases the bleeding by leaving an open wound without smearing on top; this is the reason why laser machining could result in an improved healing time. It is known that an increased pulse laser energy will inherently result in a larger ablation volume, but regarding the effect of pulse width, Beltran et al. [32] showed, in a limited width range (i.e., microsecond range, 244–388 µs) that there was no direct relation and instead proposed that the driving factor in the ablation efficiency is only the pulse energy. Additionally, Beltran et al. [33] explained how the ablation rate (i.e., ablated bone in depth per unit time) depends on both the feed speed and the number of passes of the laser beam, generally showing that for a single laser pass, the ablation rate increases while lowering the feed. However, they also suggested that a single pass results in a more uneven trench (i.e., with ripples), thereby showing that there is a non-trivial relationship between feed speed, cutting efficiency and surface quality.

Bone machining via non-conventional methods is under current investigations towards making the process a more reliable one in surgical environments. However, there is still lack of understanding on the relation between laser type (in terms of laser pulse width) and induced necrosis, as most of the aforementioned studies always employ cooling methods to the laser machining process. Additionally, most studies aiming on improving the cutting efficiency (i.e., maximising material removal rate per unit time) have been limited to a fixed pulsed width laser, thus making the processing recommendations pulse-width dependent. To bring more light into these aspects, here, the effect of pulse width of the laser beam on damage cortical bone both in terms of its surface characteristics (e.g., cracks, roughness) and necrosis is investigated. This was done for an Nd-YAG laser with pulse widths in the nanosecond, picosecond, and continuous ranges. Additionally, various feed speeds (i.e., 0.25–45 mm/s) were employed to understand the machining efficiency of each case. Histological analysis was used as a gold standard technique for assessment of the necrotic depth after machining, while metrology equipment was used for evaluation of the machined trench profiles post-machining. In this paper, the relation of laser pulse width with necrosis inducement is shown to be non-trivial, highlighting that variations of pulse width can produce both carbonisation and necrotic damage, as well as mild machining with minimum damage. It is also shown that the use of pulse widths in the picosecond range can shift the material removal mechanism from thermal to optomechanical, thereby reducing the inducement of thermal damage.

Rights and permissions

Open Access This article is licensed under a Creative Commons Attribution 4.0 International License, which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons licence, and indicate if changes were made. The images or other third party material in this article are included in the article’s Creative Commons licence, unless indicated otherwise in a credit line to the material. If material is not included in the article’s Creative Commons licence and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder. To view a copy of this licence, visit


This article is autogenerated using RSS feeds and has not been created or edited by OA JF.

Click here for Source link (