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DENTAL IMPLANTS
Introduction:
The current success and popularity of dental implants is based on the studies and findings by Branemark et al. Branemark proposed the concept of ‘osseointegration’ which he defined as ‘ a direct structural and functional connection between ordered, living bone and the surface of a load- carrying implant’. The currently accepted definition for osseointegration is ‘ contact established between normal and remodeled bone and an implant surface without the interposition of non- bone or connective tissue, at the light microscopic level. That’s why for dental implants to be an accepted treatment modality for edentulous patients, the interaction of host tissue with the underlying implant surface must be understood. Response of host tissue to dental implants is divided into two distinct but interrelated phases: clinical healing phase and functional phase.
During the clinical healing phase following placement; cellular attachment, migration and differentiation occurs. Therefore it is important to understand the fundamental characteristics like material selection and physical and chemical properties of implant surface which will affect the initial formation of host- tissue interface.
In the functional phase; bone interface remodels under the occlusal forces placed on the implant. Tissue remodelling is strongly influenced by characteristics of loading and distribution of stress at the interface.
The range of metals used for implants has become limited to cobalt- chromium(Co-Cr) alloy, commercially pure Titanium( Ti ) and it’s major biomedical alloy Ti- 6Al- 4V, as these provide good corrosion resistance and reasonable fatigue life. These alloys are much stiffer than cortical bone. Clinical implants made from both materials appear to be successful, but it is more important to determine the effects of surface morphology and texture of these materials on long- term tissue remodeling.
An approach to enhance tissue responses at dental implant interfaces has been the introduction of ceramic- like calcium-phosphate containing(CP) materials as implant devices. One of the most important uses of CP materials has been as a coating on metallic substrates. Most commonly used material to coat metallic substrates is Hydroxyapatite(HA). Relatively thin HA material is coated as thin- films to the substrate to enhance bone responses at implant sites.
Titanium alloy is more suitable than Co-Cr alloy to be coated with HA because it has less potential proximal stress shielding and bone resorption. Titanium alloy also demonstrated a 33% increase in bonding strength to the HA coating in vitro compared to Co-Cr alloy. It was suggested that this increase arouse because of the formation of a chemical bond between Ti and HA in addition to the expected mechanical locking.
The coefficient of thermal expansion is another important factor to consider, and again there is more of an agreement between titanium as opposed to Co-Cr alloy and HA. The adden advantage of a Ti alloy is the low density and good bone bonding capacity.
Structure and chemistry of HA:
Mineral phase of bone and teeth consists %60- 70 of the overall composition. This mineral phase can be described as a calcium phosphate with an apatitic structure close to that of hydroxyapatite(Ca10(PO4)6(OH)2). Hydroxyapatite occurs both geologically as a mineral and as the predominant inorganic component of bones and teeth. While geologic hydroxyapatite may occur as large, single crystals, bone and tooth mineral consists of small crystals a few angstroms in length, containing impurities such as carbonate, magnesium and acid phosphate.
HA is biocompatible and bioactive in the body. It is compatible with various types and can adhere directly to osseous, soft and muscular tissue without an intermediate layer of modified tissue. It also displays an osteoconductivity; a property of a material to encourage bone already being formed, to lie closely to or adhere ro it’s surface. This is especially useful for an implant where fast healing is required.
Despite it’s excellent properties as a biomaterial, the ceramic nature and the inherent mechanical properties of HA- specifically brittleness, poor tensile strength and poor impact resistance have restricted it’s application in many load- bearing applications. Therefore, the concept of applying HA onto metallic implants as a surface coating was developed, and the HA- coated implant combined the good strength and ductility of the metal with the excellent biocompatibility and bioactivity of the HA.
HA coatings were first introduced in the middle 1980s for improved fixation between bone and implant. Since that time, these materials have been extensively used in orthopaedic and dental implants.
Coating methods:
Various methods have been used to deposit HA coatings, such as dip coating- sintering, immersion coating, electrophoretic deposition, hot isostatic pressing(HIP), solution deposition, ion- beam sputter coating and dynamic mixing, thermal spraying techniques such as plasma spraying, flame spraying and high- velocity oxy- fuel combustion spraying.
Among these methods;
Dip coating/ high temperature sintering can degrade mechanical properties of metal implants and lead to low bond strength and impurity of HA.
Electrophretic deposition leads to nonuniform thickness of HA.
Immersion coating occurs at high temperature and results in a coating of non-HA compound mixture and very poor adherence.
With hot isostatic pressing, it’s difficult to seal borders on implants with complex shapes.
Solution- deposition is a low- temperature precipitation procedure resulting in pure, highly crystalline, firmly adherent HA coating. Maximum thickness of 20 µm limits it’s use as a primary mode of fixation, especially in dentistry.
Sputter coating is too slow and has a low deposition rate. Ca/P ratio of the coating is higher than that of synthetic HA.
So far, thermal spraying, especially the conventional atmospherical plasma- spray method appears to be the most favorable and most commonly used for clinical application. This method has a high deposition rate, good chemical and microstructure control, biocorrosion resistance, and substrate fatigue resistance of the coating. Various coating thickness can be obtained and used for complex shapes.
Figure 1. Dental implant being sprayed
Before plasma spraying takes place, the surface of the metallic implant can be textured into microstructured, macrostructured, or porous morphologies. The microstructuring can be grit blasting or beading. Grit blasting is a surface preperation method used prior to the application of plasma sprayed coatings; it alters the smoothness of the metal surface to produce a roughness of around several micrometers. This method has proven successful for implant fixation and is currently the major method for implants in clinical use. Macrosturucturing can be in the form of grooves, threads, meshes, or a deposited metal coating. Improved fixation has been demonstrated with grooved-surface implants in vivo both in the initial period and 1 year after implantation. The purpose of macrostructuring is to increase the shear strength between the substrate and the coating, and this could result in improved long- term fixation between the implant and the bone if biodegradation of the coating occurs.
The use of porous coatings is another method for implant surface treatment, and such structures have been shown to be of more benefit than a grit-blasted surface.
Critical quality specifications for HA coatings include purity, crystallinity, Ca/P ratio, microstructure, porosity, surface rougness, thickness and implant type and surface texture which also lead to different mechanical properties.
Purity and crystallinity:
The typical feedstock for HA coatings is a fully crystalline pure HA powder. It’s manufactured by chemical precipitation from a mixture of a Ca-ion-containing solution and a phosphate- containing solution. After plasma spraying, both purity and crystallinity decreases because of decomposition of HA at high temperature and rapid cooling. Low crystallinity causes high dissolution of the implant, but high crystallinity may lead to release of coating segments.
Microstructure, porosity and roughness:
The porosity of a commercially available HA coating may vary from 1% to 10% and sometimes up to 50%. When different feedstock parameters and spray parameters are employed, the original particals can become well- flatetened splats, accumulated splats, spheroidized particles, partially melted particles or remain unmelted. These different features lead to different forms of porosity and nature of porosity is very important because this controls physiochemical interactions at the implant- host interface.
Figure 2. Surface morphology of plasma- sprayed HA coatings.
(A) partially melted large particle, (B) partially melted fine particle,
(B) flattened splat, (D) accumulated splats, (E) spheroidized particle
Surface roughness of the HA coating also affects it’s dissolution and the bone apposition on the coating or bone ingrowth because the surface is directly in contact with the tooth and the body fluid. High surface roughness will increase the dissolution rate and apatite precipitation.
Thickness:
Thickness of the HA coating affects both its resorption and mechanical properties. Thicker coatings usually exhibit poorer mechanical properties. Thickness of HA coating for dental implants can be several hundred micrometers.
Mechanical Properties:
Mechanical properties are important for the long- term performance of the HA-coated implants. According to the proposed standards, the shear strength should be 22- 29 MPa and the minimum tensile strength should be 51 MPa for the HA coating. But the understanding of other most important mechanical properties like Young’s modulus, thermal stress, fracture toughness and fatigue life is still incomplete.
Dental implants coated with HA:
Dental implants include subperiosteal, transosteal and most commonly used endosseous implants. Endosseous implants can take the form of root consisting of a post, an abutment and a crown. HA is usually coated to the surface of dental posts. After careful preparation of the implant site, the post is placed into the maxilla or mandible. The gingiva covers the implant and a 3- month healing phase enables new bone to form close to the immobile implant. The implant is then opened, and the crown is mounted with an abutment. For the next 18 months, the newly formed bone remodels according to the magnitude, direction and frequency of the applied force.
Excellent clinical results were reported for HA- coated implants. HA- coated implants have higher integration rate, promote faster bone attachment, achieve direct bone bonding with higher interfacial attachment strength to bone when compared with non- coated metallic implants. They’re also osteoconductive, they can promote the growth of bone into areas that they would normally occupy. HA coatings have also suggested to have good sealing affects against the migration of polyethylene particles along the bone- implant interface, which may reduce the incidence of osteolysis and the subsequent implant failure.
HA coatings have been shown to achieve a very strong bond with living bone, in a relatively short period, even under loaded conditions and with the presence of a gap.
The process has been suggested to be initated with the dissolution of the coating soon after the implantation and can be described as:
a)parital dissolution of HA coating where calcium and phosphate ions are released from the coating which will cause a rise of the calcium and phosphate ion concentration in the local environment around the coating
b)precipitation of crystals on HA coating and ion exchange with surrounding tissues
c)formation of a carbonated calcium phosphate layer of microcrystals and macrocrystals with the incorporation of a collageneous matrix and bone growth toward the implant
d)bone remodeling in area of stress transfer: osteoclasts resorb normal bone by actively secreting hydrogen ions into the extracellular space, creating a local pH around 5 and leading to fast resorption of both carbonated HA in bone mineral and the HA coating
e)bone–implant interface is subjected to further bone ingrowth and remodeling and a biological fixation can be achieved through the bidirectional growth of a bonding layer.
Besides having lots advantages listed above, HA coatings may cause some problems in the implants, like coating resorption, wear and osteolysis.
Resorption of the coating, or coating loss is essential for the establishment of bone- implant bonding, but if it is not kept at optimal levels, complete loss of the implant may happen.
Another concern about the clinical use of HA coating is that it will lead to increased polyethylene wear and osteolysis. Osteolysis was defined as any foacl area of bone loss adjacent to the prosthesis, which is believed to be caused by the biological response to wear debris, mainly polyethylene particles, but also including metallic and HA particles. But studies made up to date have shown that HA resorption and probably fragmentation most likely do occur but do not seem to promote increased wear and osteolysis. On the other hand, becuase of it’s biocompatibility and potential circumferential bone apposition, HA coatings may prevent polyethylene wear and osteolysis.
Apart from resorption, wear and osteolysis, there are still many concerns about the use of HA coatings, especially with regard to long- term stability and resistance to oral microorganisms.
Long term stability:
HA coatings are reactive, with the potential to dissolve and lose structural integrity. Resorption and degradability of HA coatings on biological environment can lead to disintegration of the coating, resulting in the loss of both coating- substrate bond strength and implant fixation. Many studies have been done to decide if coating resorbs and compromises implant survival.
In one study, surface morphology, composition and structure at different depths of HA coating was investigated and changes occured in implant by time were characterized.
Two unused and two retrieved HA- coted titanium endosseous blade type implants were examined.
Surface morphology:
Surface examination was made by stereomicroscopic and secondary electron (SE) microscopic observation. It was seen that unused implants revealed uniform roughness of the HA- coated surfaces.
Figure 3. Surface morphology of unused HA- caoted Ti blade- type implant (stereoscopic image)
Figure 4. Surface morphology of unused HA- caoted Ti blade- type implant (SE image)
Both retrieved implants presented partial loss of HA coating, titanium substrate being visible especially in the apical and mesiodistal edges of the blades. The morphological aspects of the retrieved implants examied in this study suggested that HA loss around the apical and mesiodistal edges of the blades occured mainly by resorption, while HA flattening in the central and cervical areas was determined by wear. Different stress values in bone along the implant may determine different bone remodeling responses: higher stress areas may have more intense remodeling acitivity, so having a higher rate of resorption.
Figure 5. Surface morphology of retrieved HA- caoted Ti blade- type implant (stereoscopic image) arrows indicate attached bone fragments
Figure 6. Surface morphology of retrieved HA- caoted Ti blade- type implant (SE image) increased roughness and rounded contours of the HA surface in proximity of the apical and mesodistal edges
Section analysis:
Observation of the unused implant cross- sections revealed uneven thickness of the coating. But the retrieved implants displayed more variation of the coating thickness, due to parital loss of coating. Element analyses of the unused implant cross- secitons showed the presence of small amounts of Cl and Mg within the coatings. The relative amounts of Cl and Mg were increased in retrieved implant cross- sections. Mg has been demonstrated to compete with Ca for the same apatite lattice, while Cl competes with OH. Line profiles showed that ion exchange progressed throughout the entire coating thickness.
Figure 7. chlorine and magnesium distributions in the unused implant
Figure 8. chlorine and magnesium distributions in the retrieved implant
XRD:
The crystal structure of the coatings was investigated with an X-ray diffractometer. All sample XRD patterns revealed a basic apatite structure, proving that no major structural change occured in the retrieved implants.
FTIR:
FTIR spectra of all samples exhibited OH and PO4 absorption bands characteristic for HA. For the retrieved implants, the OH peak intensity was lower than that of the unused implants at all analyzed depths, indicating a significant loss of OH groups in the retrived implant coatings. One candidate for substitution in the OH site is Cl, and these results were consistent with element analysis.
As a result, long- term implantation did not determine major structural changes in the apatite crystal lattice, but induced compositional changes. And morphological changes observed on the implant surfaces seem to depend more on the stress values in the surrounding bone and on the implant mobility. But these changes did not seem to influence implant failure.
Also long- term follow-up studies were made with different study groups. If loss of HA coating causes late implant failure, it would be logical to assume that the yearly interval survival rate would progressively decrease with each year of follow-up. Yearly interval survival rates for the integral implant were derived from 5 different clinical trials, the survival rate remained above 90% in all 5 studies, there was no decrease in survival rate with long- term follow-up.
Figure 9. yearly interval survival rates for integral HA- coated implant for 5 different studies
In another study to understand the possibility of the existence of HA detachment and resorption, two retrieved dental implants after 12 months of loading were tested. These implants showed abutment fracture and close contact between the bone and the coating with no gap or connective tissue capsule at the interface under light microscopy.
A reduction in the coating thickness was also found, just in one area, along with the presence of some detached HA particles embedded in the newly formed bone. This suggests that the resorption or the detached HA particles would not cause any adverse problem for the long- term survival of the implant.
Infection with a pathogen:
The second major factor in implant failure in the oral environment is bacterial involvement. HA- coated dental implants may have an increased risk of infection. Specific pathogenic bacteria have been observed to preferentially bind to HA. Reports have emerged in the literature that describe evidence of bacterial invasion into the HA- metallic interface of failed dental implants. Bacterial acidic products can degrade HA coatings and facilitate further damage of the coating at the metallic-HA interface.
Figure 10. typical failed oral implant, edenmatous gingival tissue with evidence of pus present
In vitro studies have demonstrated that specific bacteria can more easily adhere to HA powder/beads than titanium powder/beads and suggested that HA coatings are more susceptible to bacterial colonization than titanium implants or natural teeth because of the surface roughness and hydrophilicity of HA. However, many clinical microbiologic studies do not seem to support this concern. Microbial colonization of HA-coated and titanium implants in a 10-month study did not show any significant difference in the development of microflora.
In a specific study, fatigue properties of HA- coated dental implants after exposure to a periodontal pathogen were studied. When HA- coated Ti alloy implant was examined, there were signs of coating disintegration. Primary failure was in HA/metal interface. Three types of organisms, namely Actinobacillus actinomycetemcomitans(Aa), Porphyromonas gingivalis(Pg) and Prevotella intermedia, have been identified as the major species associated with periodontitis. It was found that Aa could grown easily under anaerobic conditions. In this study, fatigue properties of HA-coated dental material(Ti6Al4V) before and after exposure to the periodontal pathogen Aa.
This study demonstrated that significant Calcium and phosphorus losses from the HA-coated samples in the presence of the organism Aa. Both calcium and phosphorus losses were seem to increase with the decrease of crystallinity and increase in pH.
In addition to bacterial exposure, specimens were also subjected to fatigue loading under a compression loading for 5 million cycles. Calcium loss due to fatigue was observed without any additional phosphorus loss. This observation indicated that the action of bacteria and fatigue stress cumulativelyled to the calcium loss from the coating, but not on phosphorus loss. But the reason for this observation was not clear. Both calcium and phosphorus dissolutions were higher during the bacterial culture step than those during fatigue testing, this would suggest a cumulative effect between the bacterial exposure and fatigue affecting the stability of the coating.
Do they have any advantages over uncoated Ti implants?
Besides the increased risk of infection and long term stability, the other most discussed thing about HA coated Ti implants is their advantages over uncoated Ti implants.
Several studies have been done comparing properties of HA-coated and uncoated Ti implants.
Since the initial development of HA-coated dental implant in 1984, numerous studies have demonstrated favorable or superior results for HA- coated implants as compared with uncoated titanium implants. These findings have been attrinuted to a unique interface between the HA surface and the bone.
In one long-term study, a 5-year comparison of HA- coated Ti plasma-sprayed and Ti plasma-sprayed cylinder dental implants was done. Long- term success associated with these two implant systems were investigated. Other factors considered were the gender and age of the patient, site of implant and smoking status of the patient. Loss of implant was considered as failure in this particular study. 65 patients were randomly introduced with TPS and HA- coated TPS cylinder implants. Total of 351 implants were placed in these patients. The results were evlauated statistically.
According to these results, of the first 15 implants that failed; 12 were TPS and 3 were HA-coated TPS implants.
In the preload period, TPS implants showed a higher failure rate than did HA- coated TPS implants.
After prosthetic loading, loss of TPS implants was 9.1% higher than HA- coated TPS implants.
Results have shown that age and gender were not significant in implant failure.
Site of implant has revealed different results: The implants were placed in four different sites which were posterior mandible, anterior mandible, posterior maxilla and anterior maxilla. Of these four sites, anterior maxilla had failure rate 13% higher than the other three sites.
Several implant legths were tried also: These were 8, 10, 13 and 15 mm. Among these different lengths, 10 mm implants had the highest failure rate, 17.4 % higher tha the other three.
The last evaluated effect in implant failure was smoking history of the patient. It was obvious that implant failure rate was significantly higher for smokers. Smoking history has a profound effect on the success of an implant both before and after loading.
Figure 11. comparison of smokers and non-smokers
As the result of this study, it was shown that there is little difference between the two systems, local and systemic factors appear to play a greater role in implant failure than does the surface of the implant. But it is obvious that clinical success of HA- coated implants to be greater than that of TPS implants after stage 1 surgery before prosthetic loading.
In one of the recent studies, a new titanium alloy was tried. This newly developed alloy named Ti-6Al-7Nb, exhibits better mechanical properties than pure titanium. Ti-6Al-7Nb dental implants with or without HA coating were placed in dogs. To investigate implant effects on cell attachment, proliferation, differentiation and matrix production or matrix mineralization, rat osteoblast-like cells (osteo-1 culture) were used. Osteo-1 cells were grown on disk-shaped Ti-6Al-7Nb substrates. Growth and cell viability curves were obtained by scanning elctron microscopy. Osteo-1 cells grown on plain culture dishes were used as controls (group 1).
This comparative study showed that both samples of Ti-6Al-7Nb alloys( HA- coated and uncoated) were biocompatible allowing the growth of osteo-1 cells.
Higher activity of bone formation was observed with HA-coated implants. But growth HA-coated cells was lower than uncoated cells. This may be due to physical properties of HA.
Figure 12. growth curves of ostoblast-like cells grown on plain Ti-6Al-7Nb surfaces(group 2) and HA-coated Ti-6Al-7Nb surfaces(group 3),compared to control group (group 1).
Cell viabilities in both groups were very similar. Cell death up to 15 days was high in HA-coated group but then, death rate decreased and cell viability increased.
Figure 13. Cell viability of group 1(control), 2( plain Ti-6Al-7Nb) and 3(HA- coated Ti-6Al-7Nb)
SEM analysis of osteo-1 cells have shown that when these cells were grown on plain Ti-6Al-7Nb alloy, there was considerable growth with formation of several layers of cells with few or any extracellular matrix(ECM). On the other hand, when these cells are grown on HA- coated Ti-6Al-7Nb alloy, there are fewer cell layers, however, higher amount of ECM is observed, which was composed of collagen fibrils.
Figure 14. SEM micrograph of osteo-1 cells
(A) osteo-1 cells grown on plain Ti-6Al-7Nb. Cells formed a confluent
monolayer
(B) osteo-1 cells grown on plain Ti-6Al-7Nb. Cells formed multiple layers with few ECM. The cells are flat and polygonal
(C) osteo-1 cells grown on HA-coated Ti-6Al-7Nb. Cells formed a confluent monolayer
(D) osteo-1 cells grown on HA-coated Ti-6Al-7Nb. Flat and polygonal shape with multiple cytoplasm processes of varying lengths. These cells form an ECM, mostly represented by fibrils interconnecting different cells.
This study shows two important advantages of HA-coated implants:
1- Higher activity in new bone formation is observed in HA- coated surfaces.This bone formation could be due to both the cell death and the cell differentiation observed in the HA-coated Ti-6Al-7Nb alloy samples. The cell death would be chemoattractant to macrophages elicitng chronic inflammatory reaction, hence enhancing tissue healing.
2- HA supports more rapid differentiation of cells.
Conclusion:
Plasma-sprayed HA coatings have been used as surface coatings on metallic implants in dentistry since the mid 1980s. Plasma sprayedHA- coated Titanium implants have been successfully used to reconstruct maxilla and mandible. The advantages that are sought in this application include (i) more rapid fixation and stronger bonding between the host bone and the implant, and (ii) increased uniform bone ingrowth and/or outgrowth at the bone- implant interface. HA cotaed implants also have higher integration rate, promote faster bone attachment, and achieve direct bone bonding, with higher interfacial attachment strength to bone when compared to non- coated metallic implants. HA-coated implants are osteoconductive, an inspiring fact is that HA-coating can enhance bone growth across a gap of 1 mm between the bone and the implant in both stable and unstable mechanical conditions. These coatings allow direct bonding to living tissues compared to a loosely adherent layer of fibrous tissue at the implant interface in other cementless fixation, which is especially beneficial for young active patients.
Despite these advantages and success of HA- coated implants in short and medium-term , there are still many concerns about the use of HA- coatings especially with regard to long term stability. The long-term problems can be bone resorption, wear and osteolysis, and failure because of an oral pathogen. Another controversy is the advantages of HA-coated implants over uncoated-titanium implants.
By the studies carried out up to date, the wear and osteolysis problem did not exhibit any obvious increase in HA-coated implants compared to other cemented and cemetless fixation methods. The HA coating even seemed to be able to seal the polyethylene and metal debris from entering the bone- implant interface and reduce wear and osteolysis.
Resorption of HA coating did not occur in most clinical phases, mainly cell-mediated as the result of normal bone remodeling. The loss of coating is usually followed by the growth of the new bone. Thus, the resorption rate can be optimally controlled so that the new bone can grow immediately to replace the resorbed coating, the durability of the bone-implant fixation should not be affected.
When long- term changes of HA-coated implants were examined, it appeared that long-term implantation did not determine major structural changes in the apatite crystal lattice, but induced compositional changes. But these changes did not give rise to implant failure and loss.
Effect of bacterial exposure on failure or long-term stability of HA-coated dental implants is not clear yet. However,studies on this subject revealed that bacterial exposure can cause dissolution of these implants. But subsequent fatigue loading of these samples showed more loss of coating. This suggests that there is a cumulative effect between the bacterial exposure and fatigue loading on the stability of the HA coating. But this risk is valid for all types of dental implants, it’s not a disadvantage unique to HA- coated metal implants.
When the comparison between HA-coated and uncoated Ti implants was made, it appeared that higher activity in bone formation occured when the implant surface is sprayed with HA. This can be due to higher rate of cell death, which is the chemoatractant to macrophages eliciting chronic inflammatory reaction, enhancin tissue healing. HA- coated implants also are greater than TPS implnats after stage 1 surgery before prosthetic loading. The reason for HA-coated implants to do better than Ti implants can be due to the unique interface between HA surface and the bone, which is termed ‘biointegration’.
Long- term stabilities and success of HA- coated and uncoated Ti implants does not show any significant difference but HA-coated implants achieve faster bone remodeling, faster tissue healing and more rapid differentiation of cells in the post-implantation period. Also production of extracellular matrix(ECM) has newly been discovered with HA-coated implants, which will be useful in the future studies in enhancement of bone remodeling.
In summary, the outlook on using HA coatings in dentistry, formed by thermal spray methods as functional bioactive agents to aid the healing process control in order to predetermine the precise coating chemistry and exact thickness of the HA coating will assure agreeable clinical resuts
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