Per-Ingvar Brånemark from Sweden initiated implant experiments by placing experimental implants in rabbits and dogs. He discovered that the implants could not be removed without fracturing the bone. Brånemark developed a hollow-threaded implant made of commercially pure titanium with a glass-adhered surface (Fig. 1). In his experiment, designed to demonstrate the vascularization of the implant chamber through bone marrow response, Brånemark accidentally observed that the implant chamber was the most difficult to remove from the bone. He concluded that the implant was firmly integrated into the bone tissue¹⁾. Prior to the 1980s, it was commonly accepted that soft tissue interposed between bone and implants, contradicting the prevailing concept that direct bone fixation of any metal specimen was impossible²⁾.

Figure 1. Radiographic image of Brånemark's experimental discovery of osseointegration.

The term "osseointegration" was introduced by Brånemark in 1976 and was first used in 1977 to explain implant fixation. This phenomenon was not visualized until 1982, when cutting and grinding techniques enabled the analysis of bone and metal specimens³⁾. However, Brånemark was confident that the bone fixation of implants would result in satisfactory clinical function, leading to the first oral implant surgery on a human patient⁴⁾. The direct bone fixation of oral implants was initially proposed in 1969¹⁾, but at the time, the concept of direct contact between bone and implants was not clearly visualized. The term "osseointegration" was first introduced in the title of a 1977 research paper⁴⁾.

The definition of osseointegration has evolved over time. It was initially described as "the direct contact between the load-bearing implant surface and bone at a microscopic level of resolution⁴⁾." In 1981, a more structured definition was proposed: "a direct structural and functional connection between ordered, living bone and the surface of a load-bearing implant⁵⁾." By 1990, an even more detailed definition emerged: "a sustained, structural, and functional coexistence between differentiated, appropriately remodeled biological tissue and a strictly defined and controlled synthetic component, without initiating rejection mechanisms, providing continuous and specific clinical function⁶⁾."

Although the exact mechanism of how the direct bone-implant interface forms is not yet fully defined, the concept of direct contact between the implant surface and bone is widely accepted. Osseointegration occurs in four key phases:

1.Initial Phase (Week 1): A fibrin network forms from the blood clot surrounding the implant. Platelets within the fibrin clot act as a rich source of growth factors, recruiting mesenchymal stem cells (MSC) to the implant surface. Osteoblasts migrate to the implant surface within the first week.

2.Osteoconduction Phase (1 Month): Bone morphogenetic protein (BMP), Wingless-related integration site (Wnt), and parathyroid hormone (PTH) signaling pathways are activated, leading to cell migration and osteoid formation. Immature woven bone is deposited.

3.Bone Conduction Phase (2∼3 Months): Immature woven bone transitions into mature lamellar bone.

4.Bone Adaptation Phase (3∼4 Months Onward): Bone remodeling continues in response to mechanical loading, ensuring long-term osseointegration stability.

Given the importance of implant surface characteristics in modulating cellular behavior, extensive research has been conducted on surface treatments. Implants with a surface roughness S(a) of approximately 1.5∼2.0 mm exhibit increased bone-to-implant contact compared to machined surfaces. These surfaces are typically created using various techniques such as grit-blasting, acid-etching, anodizing, and hydroxyapatite (HA) coating⁷⁾. Implant surface treatment not only increases the contact area between bone and the implant through rough surfaces but also modifies the nanoscale surface topography, enhancing gene expression related to the TGFβ-BMP signaling pathway and facilitating various biological functions^8-10).

Hydroxyapatite (HA) is a bioactive ceramic composed primarily of calcium and phosphate, known for its exceptional biocompatibility. As one of the most bioactive materials, HA has been widely utilized in tissue engineering research. The porous nature of HA mimics the inorganic composition and structure of human bone, making it more favorable for osseointegration than conventional porous coatings¹¹⁾. Even if HA coatings degrade during bone remodeling, the exposed titanium surface ensures continued osseointegration¹²⁾. The porosity of HA plays a crucial role in successful bone growth and implant fixation by enhancing cellular proliferation, adhesion, and differentiation compared to non-porous materials¹³⁾. This allows cells to adhere to the porous HA coating surface, facilitating extracellular matrix formation. The success of bone growth is directly dependent on pore size and interconnectivity¹⁴⁾. Compared to pure titanium implants, HA-coated implants exhibit a bioactive surface structure that enhances surface energy, promoting faster bone healing and improving the bone-implant interface. Porous HA coatings have been proven to positively influence healing time and implant fixation strength^15-17).

In the early stages of coating technology using HA powder deposition, some drawbacks were reported, including low crystallinity, which affected the long-term stability of the coating and led to delamination. Additionally, variations in the spraying process altered the HA substrate, and in some cases, it acted as a potential source of infection in vivo^18,19). Recently, plasma technology has been utilized to develop a post-plasma spray technique involving pressure and hydrothermal treatment, transforming HA coating into a high-crystallinity surface with a dense microstructure, minimal impurities, and strong bonding strength^19,20). This HA coating technique significantly reduces the risk of surface degradation or fracture in implants while enhancing hydrophilicity. Additionally, it prevents the biological aging of titanium caused by oxygen exposure, making it a highly recognized advancement in modern implantology (Fig. 2).

Figure 2. Evaluation of surface energy (hydrophilicity) of various implant brands using saline droplet test.

HA coating enhances bone healing by promoting bone-implant stability, accelerating osseointegration, and strengthening the bone-implant interface (Fig. 3). In plasma-sprayed HA implants, the microstructure of the HA coating is highly dense, with a porosity of 4.6%, which is lower than the internal porosity of 8.0 ± 1.0% measured in HA through SEM analysis. This indicates that the raw HA material is properly melted and integrated into the coating. Immediately after implantation, the HA layer shows slight improvements in displacement resistance and friction coefficient compared to pure titanium surfaces. However, as osseointegration progresses, the actual bone-to-implant contact (BIC) surface area and friction coefficient significantly increase, amplifying the positive effects of the osseointegration process¹⁵⁾.

Figure 3. Histological image (×100) showing direct bone contact and bone marrow space at the HA-bone interface.

For prosthetic restoration of natural teeth, the ideal crown-to-root ratio is 1:2, with a minimum of 1:1 as suggested by Shillingburg et al²²⁾. Based on this concept, early dental implants commonly exceeded 10 mm in length to increase root stability. In the early 2000s, cumulative success rates of implants shorter than 10 mm were significantly lower than those of longer implants²³⁾. However, advancements in implant surface treatment have demonstrated no significant correlation between crown-to-implant ratio and implant survival²⁴⁾. Recent studies indicate that short implants (5∼8 mm) exhibit comparable outcomes in terms of marginal bone loss, success rates, and survival rates compared to conventional implants^25,26), leading to a broader acceptance of shorter implants.

Debate continues regarding the optimal diameter of implants. Initially, wider implants were thought to provide greater bone-to-implant contact and higher primary stability, particularly for short implants²⁷⁾. However, excessive pressure on the buccal bone has been linked to increased bone resorption and gingival recession, leading to higher failure rates^28-31). Conversely, wider implants offer increased resistance to fracture^28,32). A 2019 systematic review found the highest incidence of implant fractures in 3.25 mm implants, while implants ranging from 4.0 to 5.0 mm exhibited significantly lower fracture rates³³⁾. However, these findings did not account for implant position, occlusal conditions, or prosthetic factors. Most fractures primarily occurred in 3.25 mm one-piece zirconia implants (Z-Look3), but clinicians should consider the risk of implant fracture for implants with a diameter of 4.0 mm or less. Implant fractures primarily occur at the implant neck (thread top) due to tensile and compressive stress exceeding fatigue thresholds³⁴⁾. In a 2020 study simulating peri-implantitis, where 7.5 mm of a 15 mm implant with a 3.5 mm diameter was exposed, no significant correlation was observed between fracture resistance and reducing the implant diameter to 2.6∼3.0 mm through implantoplast³⁵⁾. In terms of implant stability, no significant differences in ISQ values were found between 3.8 mm and 4.6 mm implants inserted in cancellous bone³⁶⁾.

Thus, while a minimum diameter of 4.0 mm appears necessary for fracture resistance at the implant neck, the body of the implant does not necessarily require increased thickness for this purpose. Finite element analysis has demonstrated that the introduction of a wing structure at the upper implant region effectively distributes stress, reducing peak stress at the implant-abutment connection and enhancing fracture resistance (Fig. 4).

Figure 4. Finite element analysis of an implant with a wing structure.

The Safe 3.5 implant (Withwell implant, Seoul, Republic of Korea) features a maximum body diameter of 3.5 mm, with an additional wing structure at the implant neck, increasing its diameter to 4.2 mm in the upper 0.5 mm region (Fig. 5). This implant is manufactured using Grade 5 titanium (Ti-6Al-4V), known for its superior tensile strength and yield strength. The surface is treated with HA powder blasting followed by plasma spraying at high temperatures to enhance osseointegration.

Figure 5. HA-coated implant with wing structure of Safe 3.5.

If fracture resistance is maintained, a thin-body implant offers several advantages. First, it minimizes the amount of bone removal required for implant placement, which is particularly beneficial in regions with poor bone quality. Additionally, the low taper of the implant body reduces compression on surrounding bone tissue during placement. In cases of compromised bone quality or thin cortical bone, ridge splitting can occur naturally without fracturing the bone plate. In one case, an immediate implant was placed at site #42 following extraction. After using an initial drill and a single 2.2 mm step drill, a Safe 3.5 implant (3.5 mm diameter, 8.5 mm length) was placed. The buccal bone plate remained intact and positioned properly (Fig. 6).

Figure 6. A 2.2 mm diameter step drill was used to prepare the site in the thin and low-quality mandibular anterior region, followed by the placement of the Safe 3.5 implant. The thin buccal bone plate was not removed but was naturally displaced laterally, allowing it to remain adjacent to the implant, resulting in a ridge splitting effect.

Another advantage of a thin-body implant is that it allows for repositioning if initial placement deviates from the intended site. Due to its smaller body size, implant removal causes minimal bone loss, allowing for straightforward reimplantation with minimal additional drilling. In one case, an immediate implant was planned for site #21. After using an initial drill and a 2.2 mm step drill, the implant was placed, and an autogenous tooth-derived bone graft (AutoBT, Korea Tooth Bank, Jongno-gu, Seoul) was applied. However, postoperative radiography revealed that the implant was positioned palatally within the nasopalatine duct. With patient consent, the implant was carefully removed the following day. Initial drilling was then performed again, and the implant was placed correctly using a 3.5 mm step drill (Fig. 7).

Figure 7. The implant, which had initially deviated palatally during placement, was removed and successfully repositioned in the desired location using a 3.5 mm step drill.

Another notable advantage of the wing structure in the Safe 3.5 implant is the tenting effect. In addition to the osteoconductive properties of the HA coating, the wing structure prevents soft tissue infiltration from the upper region, thereby acting as a scaffold to maintain space (Fig. 8).

Figure 8. Observation of osteoconduction beneath the wing structure in a rabbit model (available at http://www.wwimplant.co.kr/sub/news_4.php?idx=4).

The evolution of osseointegration and implant technology began with the discovery that titanium implants could achieve direct bone anchorage. Among various surface modifications developed to enhance osseointegration, HA-coated implants have played a crucial role in promoting bone healing, improving implant stability, and accelerating bone integration. The porosity of HA enhances cellular proliferation and bone growth, while advanced plasma coating techniques have improved adhesion strength and hydrophilicity.

Research on implant length and diameter has demonstrated that, contrary to the crown-to-root ratio concept in natural dentition, shorter implants (≤8 mm) can achieve long-term stability. Moreover, wider implants do not necessarily offer superior biomechanical benefits, as fractures predominantly occur at the implant neck rather than the body. The Safe 3.5 implant, with its thin body and wide-wing structure, offers several potential biomechanical advantages. Structural analysis and long-term clinical studies are required to evaluate the efficacy and stability of this novel wing-type implant design.